Production of tissue engineered digits and limbs

ABSTRACT

The invention pertains to methods of producing artificial composite tissue constructs that permit coordinated motion. Biocompatable structural matrices having sufficient rigidity to provide structural support for cartilage-forming cells and bone-forming cells are used. Biocompatable flexible matrices seeded with muscle cells are joined to the structural matrices to produce artificial composite tissue constructs that are capable of coordinated motion.

RELATED APPLICATIONS

The present application claims priority to U.S. Provisional ApplicationSer. No. 60/663,458, filed Mar. 18, 2005 and U.S. ProvisionalApplication No. 60/660,832, filed Mar. 11, 2005, the contents of whichare hereby incorporated by reference in their entirety.

BACKGROUND OF THE INVENTION

The technical field of this invention relates to artificial tissueconstructs and method of creating such constructs.

Extremity injuries arising from battlefield trauma, which have becomemore prevalent in the past few years, have resulted in an amputationrate of approximately 29 percent. In addition to battlefield trauma,accidents caused by automobile collisions, falls, operation of machineryand heavy loading, also damage or break bones and can injure connectiveand interstitial tissues. Large limb amputation, as well as loss offingers and toes leads to a significant impact on the standard of lifefor the patient. Finger amputations are particularly disabling due tothe highly functional nature of the human hand and the day-to-daydependence on that functionality.

Conventional approaches to restore function to these patients haveincluded whole hand transplants, prosthetics, as well as toe-to-fingerand finger-to-thumb replantation. These techniques aim to restore handfunction, and functional prehension by the restoration of an opposabledigit, which is key to the patient's recovery. Unfortunately, theseapproaches have resulted in sub-optimal outcomes and limited functionalrestoration.

Accordingly, a need exists for alternative methods of replacing orrepairing missing and damaged digits and limbs.

SUMMARY OF THE INVENTION

The present invention provides methods and compositions for preparingtissue engineered functional digits and limbs. This is accomplished bygenerating composite tissues (i.e., tissue containing bone, muscle,tendon, ligament, nerve, blood vessel), grown from cells isolated fromthe patient or a donor, that provide coordinated motion. In addition,“intelligent” scaffold systems are disclosed that support clinicallyrelevant tissue dimensions and restoration of digit and limb function.This is particularly important for creating limbs because theseintelligent scaffolds address the increased biological demands ofengineering large tissues, especially in a perfusion deficientenvironment. These intelligent scaffold system are capable of releasingangiogenic and neurogenic factors to enhance tissue maturation andfunctionality. The diffusional limitations can be overcome byincorporation of vascular endothelial growth factor (VEGF) whichenhances vascularisation in the muscle component of the composite tissueand also encourages recruitment of other cells. In addition, nervegrowth can be stimulated by nerve growth factor (NGF), a potent axonalguiding mediator that promotes reinnervation of tissues.

The invention involves seeding multiple cell types on a scaffold matrixsystem that is capable of maturing into a composite tissue comprising amuscle-bone composite, or muscle-cartilage composite. The scaffoldmatrix system involves using a structural matrix having sufficientrigidity to provide structural support for bone and cartilage and aflexible matrix that is able to contract and relax with the muscletissue formed thereon. The muscle tissue provides elasticity andcontractability. The structural matrix and the flexible matrix arejoined to create an artificial composite tissue construct capable ofmoving like native digits and limbs.

Accordingly, the invention pertains to a method of producing anartificial composite tissue construct permitting coordinated motion. Themethod involves using a first biocompatable structural matrix havingsufficient rigidity to provide structural support, and seeding the firstmatrix with an isolated population of cells selected from the groupconsisting of cartilage-forming cells, bone-forming cells andcombinations thereof. A second biocompatable flexible matrix is seededwith an isolated population of muscle progenitor cells (MPCs). Thestructural matrix and the flexible matrix are joined to produce anartificial composite tissue construct capable of coordinated motion.

More than one structural matrix can be used. Two separate structuralmatrices can be joined together by the flexible matrix to provide aflexible linkage between the two matrices. The tissue construct can bejoined to the subject by joining the flexible matrix to the structuralmatrix and the natural bone structure.

The structural matrix can be an electrospun substrate, a decellularizedsubstrate, or a synthetic polymer substrate with rigid properties, andthe flexible matrix can be selected a decellularized submucosa substrate(e.g., a decellularized bladder submucosa) or an electrospun substrate.

The invention also discloses a biocompatible composite scaffoldingsystem capable of providing structural support for engineered bonetissue comprising a biodegradable synthetic polymer and naturallyderived collagen matrix. The biocompatible composite scaffolds of thepresent invention are non-toxic, easily fabricated and providestructural features that enhance the formation of bone tissue fortherapeutic regeneration. The synthetic polymer can be selected from thegroup comprising poly(lactide-co-glycolide) (PLGA), poly(lactide) (PLA),poly(glycolic acid) (PGA), poly(caprolactone), polycarbonates,polyamides, polyanhydrides, polyamino acids, polyortho esters,polyacetals, polycyanoacrylates, degradable polyurethanes,hydroxyapatite (HA), tricalcium phosphate (TCP), and calcium sulfate.The naturally derived collagen matrix is submucosa, such as bladdersubmucosa (BSM). The composite scaffolding system can be fabricatedusing a solvent casting/particulate leaching process. The biocompatiblecomposite scaffolding system preferably has a substantially uniformporous structures having an average pore diameter ranging from about 50to about 250 μm, preferably from about 90 to about 150 μm, and mostpreferably from about 110 to about 130 μm. The porosity of thebiocompatible composite scaffolding is greater than about 50%,preferably greater than about 80%, more preferably greater than about90%, and most preferably greater than about 94%.

In another aspect, the invention provides a method of fabricating abiocompatible composite scaffolding system capable of providingstructural support for regenerated bone tissue comprising selecting apore size, obtaining porogens of the selected size, adding the porogensto a solution containing a synthetic polymer and submucosa and mixing todistribute the porogens and form a composite scaffold, drying thecomposite scaffold to remove residual solvent, and removing the porogensfrom the composite scaffold. The method can further include sievingporogens through a sieve to obtain the selected size porogens. The stepof removing the porogens from the composite scaffold can compriseimmersion in water.

Methods of joining the structural matrix and the flexible matrixinclude, but are not limited to, joining technique such as suturing,heating, and gluing with biological glue.

The structural matrix and the flexible matrix may further comprise abiological agent selected from the group consisting of nutrients, growthfactors, cytokines, extracellular matrix components, inducers ofdifferentiation, products of secretion, immunomodulators, proteins,antibodies, nucleic acids molecules, carbohydrates, andbiologically-active compounds which enhance or allow growth of thecellular network or nerve fibers.

In one embodiment, the biological agent is a growth factor selected fromthe group consisting of transforming growth factor-alpha (TGF-α),transforming growth factor-beta (TGF-β), platelet-derived growth factor(PDGF), fibroblast growth factor (FGF), nerve growth factor (NGF), brainderived neurotrophic factor, cartilage derived factor, bone growthfactor (BGF), basic fibroblast growth factor, insulin-like growth factor(IGF), vascular endothelial growth factor (VEGF), granulocyte colonystimulating factor (G-CSF), hepatocyte growth factor, glial neurotrophicgrowth factor (GDNF), stem cell factor (SCF), keratinocyte growth factor(KGF), and skeletal growth factor.

Examples of bone-forming cells that be seeded on the structural matrixinclude, but are not limited to, osteogenic cells, osteoblasts,osteocytes, osteoclasts, and bone-lining cells. Examples ofcartilage-forming cells include, but are not limited to, chondrocytesand chondroblasts.

In another aspect, the invention pertains to a method for treating asubject with a limb or digit disorder. The method involves providing anartificial composite tissue construct having coordinated motion. Theartificial composite tissue comprises a first biocompatable structuralmatrix having sufficient rigidity to provide structural support that isseeded with an isolated population of cells selected from the groupconsisting of cartilage-forming cells, bone-forming cells andcombinations thereof, and a second biocompatable flexible matrix that isseeded with an isolated population of muscle progenitor cells (MPCs).The structural matrix and the flexible matrix are joined together suchthat the artificial composite tissue has coordinated motion. Thisartificial composite tissue is attached to a subject by joining theflexible matrix to the structural matrix and the natural bone structureof the subject such that there is a flexible linkage between thestructural matrix and the bone structure. The improvement in the limb ordigit disorder can then be monitored by measuring parameters such asmobility of the digit or limb, the extensor and flexor function, themotor function, the sensory function, the contractile response, and thetensile response.

In yet another aspect, the invention pertains to a muscle-cartilagecomposite construct or a muscle-bone composite construct made by themethod of the invention.

DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic of an electrospin apparatus;

FIG. 2 is a schematic of the conjugation of heparin on quantum dots;

FIG. 3 is an electrospin nanofiber;

FIG. 4 is graph of pressure-diameter curves of vascular graft scaffolds;

FIG. 5A is a graph of axial and circumferential stress-strain data fromuniaxial testing of two decellularized constructs;

FIG. 5B is a graph of axial and circumferential stress-strain data fromuniaxial testing of an electrospun vessel;

FIG. 6A is a graph of cell viability of endothelial cells cultured onfour matrices;

FIG. 6B is a graph of mitochondrial metabolic activity of endothelialcells cultured on four matrices;

FIG. 7A is a graph of cell viability of endothelial cells cultured onfive matrices;

FIG. 7B is a graph of mitochondrial metabolic activity of endothelialcells cultured on five matrices;

FIG. 8 is a graph of microcapsules containing heparin which show heparinrelease upon near infra-red irradiation;

FIG. 9 is a graph showing the quantification of remaining heparin fromretrieved vascular scaffolds.

FIG. 10A is a schematic showing that a functional limb requires supportand locomotion;

FIG. 10B is a schematic of a composite scaffold system;

FIG. 10C is a schematic of the artificial composite tissue construct ofthe invention permitting coordinated motion;

FIG. 11 is a graph showing that nerve growth factor increased the numberof regenerated axons compared to the controls without nerve growthfactors;

FIG. 12 is a schematic of an artificial finger;

FIG. 13A is a scanning electron microscope (SEM) cross-sectionalobservation of the BSM-PLGA composite scaffold of the present invention;

FIG. 13B is an SEM cross-sectional observation of PLGA scaffold;

FIG. 13C is an SEM cross-sectional observation of Collagraft scaffold;

FIG. 14A is a bar graph showing similar cell survival of human embryonicstem cells (hES) cells and bovine osteocytes (bOBs) cultured on variousbone scaffolds at 7 days as measured by neutral red assay;

FIG. 14B is a bar graph showing similar mitochondrial metabolic activityof hES cells and bOBs cultured on various bone scaffolds at 7 days asmeasured by MTT assay;

FIG. 15 is a bar graph comparing the percent cell adhesion on variousscaffolds with hES and bOSs cells.;

FIG. 16 is a graph comparing cell proliferation of the BSM-PLGAcomposite scaffold, PLGA scaffold and Collagraft up to 8 days culture,as determined by MTT assay.

DETAILED DESCRIPTION

So that the invention may more readily be understood, certain terms arefirst defined:

The term “attach” or “attaches” as used herein refers to cells thatadhere directly or indirectly to a substrate as well as to cells thatadhere to other cells.

The phrase “biocompatible substrate” as used herein refers to a materialthat is suitable for implantation into a subject onto which a cellpopulation can be deposited. A biocompatible substrate does not causetoxic or injurious effects once implanted in the subject. In oneembodiment, the biocompatible substrate is a polymer with a surface thatcan be shaped into the desired structure that requires repairing orreplacing. The polymer can also be shaped into a part of an structurethat requires repairing or replacing. The biocompatible substrateprovides the supportive framework that allows cells to attach to it, andgrow on it. Cultured populations of cells can then be grown on thebiocompatible substrate, which provides the appropriate interstitialdistances required for cell-cell interaction.

The term “subject” as used herein is intended to include livingorganisms in which an immune response is elicited. Preferred subjectsare mammals. Examples of subjects include but are not limited to,humans, monkeys, dogs, cats, mice, rates, cows, horses, pigs, goats andsheep.

The term “decellularized” or “decellularization” as used herein refersto a biostructure (e.g., an organ, or part of an organ), from which thecellular and tissue content has been removed leaving behind an intactacellular infra-structure. Organs such as the kidney are composed ofvarious specialized tissues. The specialized tissue structures of anorgan, or parenchyma, provide the specific function associated with theorgan. The supporting fibrous network of the organ is the stroma. Mostorgans have a stromal framework composed of unspecialized connectingtissue which supports the specialized tissue. The process ofdecellularization removes the specialized tissue, leaving behind thecomplex three-dimensional network of connective tissue. The connectivetissue infra-structure is primarily composed of collagen. Thedecellularized structure provides a biocompatible substrate onto whichdifferent cell populations can be infused. Decellularized biostructurescan be rigid, or semi-rigid, having an ability to alter their shapes.Examples of decellularized organs useful in the present inventioninclude, but are not limited to, the heart, kidney, liver, pancreas,spleen, bladder, ureter and urethra.

The phrase “three-dimensional scaffold” as used herein refers to theresidual infra-structure formed when a natural biostructure, e.g. anorgan, is decellularized. This complex, three-dimensional, scaffoldprovides the supportive framework that allows cells to attach to it, andgrow on it. Cultured populations of cells can then be grown on thethree-dimensional scaffold, which provides the exact interstitialdistances required for cell-cell interaction. This provides areconstructed organ that resembles the native in vivo organ. Thisthree-dimensional scaffold can be perfused with a population of culturedcells, e.g., endothelial cells, which grow and develop to provide anendothelial tissue layer capable of supporting growth and development ofat least one additional cultured cell population.

The term “natural biostructure” as used herein refers to a biologicalarrangement found within a subject, for example, organs, that includebut are not limited, heart, kidney, liver, pancreas, spleen, bladder,ureter and urethra. The term “natural biostructure” is also intended toinclude parts of biostructures, for example parts of organs, forexample, the renal artery of a kidney.

The terms “electrospinning” or “electrospun,” as used herein withreference to electrospun materials refers to any method where materialsare streamed, sprayed, sputtered, dripped, or otherwise transported inthe presence of an electric field. The electrospun material can bedeposited from the direction of a charged container towards a groundedtarget, or from a grounded container in the direction of a chargedtarget. In particular, the term “electrospinning” means a process inwhich fibers are formed from a charged solution comprising at least onenatural biological material, at least one synthetic polymer material, ora combination thereof by streaming the electrically charged solutionthrough an opening or orifice towards a grounded target.

A natural biological material can be a naturally occurring organicmaterial including any material naturally found in the body of a mammal,plant, or other organism. A synthetic polymer material can be anymaterial prepared through a method of artificial synthesis, processing,or manufacture. Preferably the synthetic materials is a biologicallycompatible material. The natural or synthetic materials are also thosethat are capable of being charged under an electric field.

The terms “solution” and “fluid” as used herein describe a liquid thatis capable of being charged and which comprises at least one naturalmaterial, at least one synthetic material, or a combination thereof. Ina preferred embodiment, the fluid comprises at least one type ofcollagen, an additional natural material such as at least one type ofelastin and at least one synthetic polymer, e.g., poly-L glycolic acid(PLGA).

The term “co-polymer” as used herein is intended to encompassco-polymers, ter-polymers, and higher order multiple polymercompositions formed by block, graph or random combination of polymericcomponents.

The term “intelligent scaffold” as used herein refers to a biocompatiblesubstrate that is able to recruit cells, encourage cell growth andmaturation, and facilitate the formation of a vascular support (e.g.blood vessels) and innervation (nerve formation) via controlled releaseof factors, such as vascular endothelial growth factor (VEGF) and nervegrowth factor (NGF). The intelligent scaffold possesses theultrastructural, biochemical and biological characteristics required forcell attachment, neovascularization, innervation and tissue maturation,even in an environment with a limited blood supply and is able to meetthe metabolic demands of the growing cells.

The term “coordinated motion” as used herein refers to an artificialcomposite tissue construct that is made up of one or more structuralmatrices attached together by a flexible matrix such that to providesynchronized or harmonized movement and locomotion The coordinatedmotion mimics the motion of native digits and limbs. For example, anative finger or thumb is able to bend and flex, and likewise, theartificial composite tissue construct is able to bend and flex.

I. Limbs

In one aspect, the invention pertains to creating a tissue engineereddigits and limbs. The limb system performs numerous tasks, includingstructural support and locomotion. It is composed of many tissues,including the bone, cartilage, tendon, muscle, nerve and blood vessels.Composite tissues that provide coordinated motion can be generated usingmultiple tissue types that are engineered to restore a portion of adigit.

In spite of the diversity of the form and function of various limbs, alllimbs have substantial physiological similarities to each other. Forexample, limbs are typically composed of the same essential materials,regardless of specialized function or shape. A significant portion ofthe limb is composed of bone and muscle. Bone primarily has organicmaterial, mostly protein-associated glycosaminoglycans and especiallycollagen, a protein commonly found in connective tissues and inextracellular matrixes. About half of the bone mass is mineral, and themost common bone-associated mineral is a calcium compound that closelyresembles hydroxyapetite.

Bone typically has bone-producing cell types such as osteogenic cells,osteoblasts, and osteoclasts, for example. Another example ofphysiological similarities shared by various types of bones isvascularization. Blood vessels penetrate and permeate the bone through aseries of channels and canals. These channels and canals can vary inlocation, length, and diameter. Examples of such channels include thecanaliculi, haversian canals, osteons, Volkmann's canals, and others.These canals bring blood, nutrients, and other vital factors to andcarry away waste from the multitude of cells that live in and that formthe bone. One important example of such a bone-dwelling cell is thepluripotent stem cell that can differentiate into and form any of thecells of the blood or immune system. The blood and immune cellsthemselves also reside, at least for a time, in the bone. Such cellsinclude macrophages, neutrophils, B cells, various T cells, eosinophils,basophils, megakaryocytes (the progenitor of platelets), and red bloodcells. There are also varieties of bone-forming cells such as theosteoclasts, osteoblasts, and osteogenic cells mentioned above that livein the bone. Attached to the bones may be various tendons and otherconnective tissues that cooperate with the bones and any connectedmuscles to establish the articulation of movement or to secure tissuesin place.

Yet another example of shared bone physiology is the specializedstructures located at one or both ends of some bones. These specializedstructures are adapted to hold a bone in a joint, while at the same timegiving some degree of freedom and relative motion to the jointed bones.Examples of this type of arrangement include the bones that form theelbow, knee, shoulder, finger, ankle, foot, vertebral, and similarlymovable joints. In addition to these and other physiologicalsimilarities, bones also can share developmental similarities. Forexample, the process of either intramembraneous ossification orendochondral ossification forms the bones, wherein the latter processinvolves a cartilaginous intermediate that is calcified to form theresulting bone tissue. Fibroblast cells typically lay down a network ofcollagen fibrils, upon which are deposited the calcium crystals in amineral form approximating hydroxyapetite.

As can be understood by those skilled in the art, the complexity of bonefunctions and structures establish a skeletal system having significantcapabilities. Bone tissue is not only a mechanical structure, but isalso a complex and diverse combination of living, dynamic, andcontinuously regenerating tissues.

The anatomy of digits and limbs can readily be determined from forexample, Gray's Anatomy, The Anatomical Basis of Medicine and Surgery,39th Edition (2005). The methods and compositions of the invention areused for the reconstruction of artificial digits and limbs that mimicthe native in vivo digit and limits structure. In particular, theinvention pertains to generating digits and limbs of the hand. The handcontains three main groups of bones, the carpal bones, metacarpal bones,and the phalanges. Additionally, the hand also contains some small bonystructures termed sesamoid bones. The carpel bones consist of thescaphoid, lunate, triquetrum, pisiform, trapezium, trapezoid, capitate,and hamate bones. The carpal bones join each other by planer-typejoints. The carpal tunnel transmits the median nerve and the tendons

The carpal bones joins to the metacarpal bones to form thecarpometacarpal joint. The hand has five metacarpal bones. The firstmetacarpal bone constitutes the skeleton of the thumb. The other fourmetacarpal bones contact with the trapezoid, capitate and hamate, andlateral-medial surfaces of metacarpal bones. The heads of the metacarpalbones, which form the knuckles, articulate with the proximal phalanges.The hand has 14 phalanges. Each finger contains three phalanges, buteach thumb only has two.

The hand also contains eight tendons of the flexor digitorumsuperficialis and profundus, the tendons of the flexor pollicis longus,and the flexor carpi radialis which pass throughout the carpal tunnel,and reach the carpal bones or fingers. Muscles of the hand occupy thespace between the metacarpals, such as the abductor, oppenens, andflexor muscles. The hand also has a complex and rich vascular networkprovided by the radial and ulnar arteries.

In addition, the hand contains radial nerves that provides sensoryfeedback, median nerves that provide motor and sensitive innervation tothe hand, and ulnar nerves, that has sensory and motor branches similarto the median nerve. The deep branch of the ulnar nerve innervates themuscles of the hand.

Although the hand is described in detail, the methods and compositionsof the invention can be used to create any limb using the samemethodology. The use of “intelligent scaffolds” further enhances theability to create larger limbs because these scaffolds encourage theattraction of cells to the scaffold and also provide factors such asVEGF and NGF, that induce cell growth and innervation on the scaffold.

II. Electrospun Matrices

The invention pertains to methods and compositions for producing andusing electrospun matrices. The process of electrospinning generallyinvolves the creation of an electrical field at the surface of a liquid.The resulting electrical forces create a jet of liquid which carrieselectrical charge. The liquid jets may be attracted to otherelectrically charged objects at a suitable electrical potential. As thejet of liquid elongates and travels, it will harden and dry. Thehardening and drying of the elongated jet of liquid may be caused bycooling of the liquid, i.e., where the liquid is normally a solid atroom temperature; evaporation of a solvent, e.g., by dehydration,(physically induced hardening); or by a curing mechanism (chemicallyinduced hardening). The produced fibers are collected on a suitablylocated, oppositely charged target substrate.

The electrospinning apparatus includes an electrodepositing mechanismand a target substrate. The electrodepositing mechanism includes atleast one container to hold the solution that is to be electrospun. Thecontainer has at least one orifice or nozzle to allow the streaming ofthe solution from the container. If there are multiple containers, aplurality of nozzles may be used.

One or more pumps (e.g., a syringe pump) used in connection with thecontainer can be used to control the flow of solution streaming from thecontainer through the nozzle. The pump can be programmed to increase ordecrease the flow at different points during electrospinning.

The electrospinning occurs due to the presence of a charge in either theorifices or the target, while the other is grounded. In someembodiments, the nozzle or orifice is charged and the target isgrounded. Those of skill in the electrospinning arts will recognize thatthe nozzle and solution can be grounded and the target can beelectrically charged.

The target can also be specifically charged or grounded along apreselected pattern so that the solution streamed from the orifice isdirected into specific directions. The electric field can be controlledby a microprocessor to create an electrospun matrix having a desiredgeometry. The target and the nozzle or nozzles can be engineered to bemovable with respect to each other thereby allowing additional controlover the geometry of the electrospun matrix to be formed. The entireprocess can be controlled by a microprocessor that is programmed withspecific parameters that will obtain a specific preselected electrospunmatrix.

In embodiments in which two materials combine to form a third material,the solutions containing these components can be mixed togetherimmediately before they are streamed from an orifice in theelectrospinning procedure. In this way, the third material formsliterally as the microfibers in the electrospinning process.

While the following is a description of a preferred method, otherprotocols can be followed to achieve the same result. In FIG. 1, acontainer 10, (e.g., a syringe or micropipette), with an orifice ornozzle 12 (e.g. a Taylor cone), is filled with a solution with at leastone natural material, and at least one synthetic material 14. Thecontainer 10 is suspended opposite a grounded target 16, such as a metalground screen. A fine wire 18 is placed in the solution to charge thesolution in the container to a high voltage. At a specific voltagedetermined for each solution, the solution in the container nozzle isdirected towards the grounded target. The single jet stream 20 ofmaterials forms a splayed jet 22, upon reaching the grounded target,e.g., a rapidly rotating mandrel. The splayed jet collects and dries toform a three-dimensional, ultra thin, interconnected matrix ofelectrospun fibers. In some embodiments, a plurality of containers canbe used with each of the containers holding a different compound.

Minimal electrical current is involved in the electrospinning process,therefore the process does not denature the materials that form theelectrospun matrix, because the current causes little or no temperatureincrease in the solutions during the procedure.

The electrospinning process can be manipulated to meet the specificrequirements for any given application of the electrospun matrix. In oneembodiment, a syringe can be mounted on a frame that moves in the x, yand z planes with respect to the grounded substrate. In anotherembodiment, a syringe can be mounted around a grounded substrate, forinstance a tubular mandrel. In this way, the materials that form thematrix streamed from the a syringe can be specifically aimed orpatterned. Although the micropipette can be moved manually, the frameonto which the a syringe is mounted can also be controlled by amicroprocessor and a motor that allows the pattern of streaming to bepredetermined. Such microprocessors and motors are known to one ofordinary skill in the art, for example matrix fibers can be oriented ina specific direction, they can be layered, or they can be programmed tobe completely random and not oriented.

The degree of branching can be varied by many factors including, but notlimited to, voltage (for example ranging from about 0 to 30,000 volts),distance from a syringe tip to the substrate (for example from 1-100 cm,0-40 cm, and 1-10 cm), the speed of rotation, the shape of the mandrel,the relative position of the a syringe tip and target (i.e. in front of,above, below, aside etc.), and the diameter of a syringe tip(approximately 0-2 mm), and the concentration and ratios of compoundsthat form the electrospun matrix. Other parameters which are importantinclude those affecting evaporation of solvents such as temperature,pressure, humidity. The molecular weight of the polymer improves itsability to entangle and form fibers, and polymers with the molecularweight of 100 kDa generally performed. Those skilled in the art willrecognize that these and other parameters can be varied to formelectrospun materials with characteristics that are particularly adaptedfor specific applications.

The geometry of the grounded target can be modified to produce a desiredmatrix. By varying the ground geometry, for instance having a planar orlinear or multiple points ground, the direction of the streamingmaterials can be varied and customized to a particular application. Forinstance, a grounded target comprising a series of parallel lines can beused to orient electrospun materials in a specific direction. Thegrounded target can be a cylindrical mandrel whereby a tubular matrix isformed. The ground can be variable surface that can be controlled by amicroprocessor that dictates a specific ground geometry that isprogrammed into it. Alternatively, the ground can be mounted on a framethat moves in the x, y, and z planes with respect to a stationarycontainer, e.g., a syringe or micropipette tip.

Electrospinning allows great flexibility and allows for customizing theconstruct to virtually any shape needed. In shaping matrices, portionsof the matrix may be sealed to one another by, for example, heatsealing, chemical sealing, and application of mechanical pressure or acombination thereof. The electrospun compositions may be shaped intoshapes such as a skin patch, an intraperitoneal implant, a subdermalimplant, the interior lining of a stent, a cardiovascular valve, atendon, a ligament, a muscle implant, a nerve guide and the like.

The electrospinning process can also be modified for example by (i)using mixed solutions (for example, materials along with substances suchas cells, growth factors, or both) in the same container to producefibers composed of electrospun compounds as well as one or moresubstances to produce a “blend,” and (ii) applying agents such as Teflononto the target to facilitate the removal of electrospun compounds fromthe target (i.e. make the matrix more slippery so that the electrospunmatrix does not stick to the target).

The various properties of the electrospun materials can be adjusted inaccordance with the needs and specifications of the cells to besuspended and grown within them. The porosity, for instance, can bevaried in accordance with the method of making the electrospun materialsor matrix. Electrospinning a particular matrix, for instance, can bevaried by fiber size and density. If the cells to be grown in the matrixrequire a high nutrient flow and waste expulsion, then a loose matrixcan be created. On the other hand, if the tissue to be made requires adense environment, then a dense matrix can be designed. Porosity can bemanipulated by mixing salts or other extractable agents. Removing thesalt will leave holes of defined sizes in the matrix.

One embodiment for appropriate conditions for electrospinning a matrixis as follows. For electrospinning a matrix by combining 45% collagen I,15% elastin and 40% PLGA, the appropriate approximate ranges are:voltage 0-30,000 volts (10-100 kV potential preferably 15-30 kV); pH 7.0to 8.0; temperature 20 to 40° C., e.g., room temperature of 25° C.; andthe distance from the container to the grounded plate can range fromabout 1 cm to about 100 cm, preferably about 1 cm to 10 cm. In additionto depositing the charged fibers on the grounded plate, the fibers canbe deposited onto another substrate such as a stainless steel mandrel.The mandrel can be rotated at 20-1000 rpm, preferably about 300-700 rpm.

Examples of naturally occurring materials include, but are not limitedto, amino acids, peptides, denatured peptides such as gelatin fromdenatured collagen, polypeptides, proteins, carbohydrates, lipids,nucleic acids, glycoproteins, lipoproteins, glycolipids,glycosaminoglycans, and proteoglycans. In a preferred embodiment, thematerials compound is an extracellular matrix material, including butnot limited to collagen, fibrin, elastin, laminin, fibronectin,hyaluronic acid, chondroitin 4-sulfate, chondroitin 6-sulfate, dermatansulfate, heparin sulfate, heparin, and keratan sulfate, andproteoglycans. These materials may be isolated from humans or otheranimals or cells. A preferred natural compound is collagen. Examples ofcollagen include, but are not limited to collagen I, collagen II,collagen III, collagen IV, collagen V, collagen VI, collagen VII,collagen VIII, collagen IX, and collagen X. Another preferred naturalcompound is elastin. Elastin fibers are responsible for the elasticproperties of several tissues. Elastin is found, for example, in skin,blood vessels, and tissues of the lung where it imparts strength,elasticity and flexibility.

One class of synthetic polymer materials are biocompatible syntheticpolymers. Such polymers include, but are not limited to,poly(urethanes), poly(siloxanes) or silicones, poly(ethylene),poly(vinyl pyrrolidone), poly(2-hydroxy ethyl methacrylate),poly(N-vinyl pyrrolidone), poly(methyl methacrylate), poly(vinylalcohol) (PVA), poly(acrylic acid), poly(vinyl acetate), polyacrylamide,poly(ethylene-co-vinyl acetate), poly(ethylene glycol), poly(methacrylicacid), polylactic acid (PLA), polyglycolic acids (PGA),poly(lactide-co-glycolides) (PLGA), nylons, polyamides, polyanhydrides,poly(ethylene-co-vinyl alcohol) (EVOH), polycaprolactone, poly(vinylacetate), polyvinylhydroxide, poly(ethylene oxide) (PEO) andpolyorthoesters or any other similar synthetic polymers that may bedeveloped that are biologically compatible. A preferred syntheticpolymer is PGLA.

In matrices composed of electrospun elastin (for elasticity),electrospun collagen (to promote cell infiltration and lend mechanicalintegrity), and other components, such as PLGA, PGA, PLA, PEO, PVA, orother blends, the relative ratio of the different components in thematrix can be tailored to specific applications (e.g. more elastin, lesscollagen depending on the tissue to be engineered).

Electrospun matrices can be formed of electrospun fibers of syntheticpolymers that are biologically compatible. The term “biologicallycompatible” includes copolymers and blends, and any other combinationsof the forgoing either together or with other polymers. The use of thesepolymers will depend on given applications and specifications required.A more detailed discussion of these polymers and types of polymers isset forth in Brannon-Peppas, Lisa, “Polymers in Controlled DrugDelivery,” Medical Plastics and Biomaterials, November 1997, which isincorporated herein by reference.

When both natural and synthetic materials are used in an electrospunmatrix, the natural material component can range from about 5 percent toabout 95 percent, preferably from about 25 percent to about 75 percentby weight. The synthetic material component can range from about 5percent to about 95 percent, preferably from about 25 percent to about75 percent by weight. In certain embodiments, both collagen and elastincan be included as natural material components, preferably with apredominance of collagen, e.g., greater than 40 percent of the naturalmaterial component. Ratios of collagen, elastin, and PLGA may betailored to fit the application: for instances, normal levels ofcollagen and elastin vary from the more elastic vessels closer to theheart to less compliant vessels further from the heart. A vessel such asthe aorta would have greater elastin content than a distal vessel. Thepercentages of collagen I, elastin, and other collagens (collagen IIIfor blood vessels or collagen II, for instance, for cartilage) may bewhatever is desired, as long as the molecular weight of these collagensis sufficient to form fibers in the electrospinning process. Ratios ofcollagen I may be from 40% to 80%, or 50%-100%. Elastin may also be usedin higher ratios from 5% to 50%. PLGA or another synthetic biodegradablepolymer may be used as desired in ratios from 5 to 80%. For a completelybiological substrate, synthetic polymers may be omitted completely andonly biological polymers may be used.

The compounds to be electrospun can be present in the solution at anyconcentration that will allow electrospinning. In one embodiment, thecompounds may be electrospun are present in the solution atconcentrations between 0 and about 1.000 g/ml. In another embodiment,the compounds to be electrospun are present in the solution at totalsolution concentrations between 10-15 w/v% (100-150 mg/ml or 0-0.1 g/L).

The compounds can be dissolved in any solvent that allows delivery ofthe compound to the orifice, tip of a syringe, under conditions that thecompound is electrospun. Solvents useful for dissolving or suspending amaterial or a substance will depend on the compound. Electrospinningtechniques often require more specific solvent conditions. For example,collagen can be electrodeposited as a solution or suspension in water,2,2,2-trifluoroethanol, 1,1,1,3,3,3-hexafluoro-2-propanol (also known ashexafluoroisopropanol or HFIP), or combinations thereof. Fibrin monomercan be electrodeposited or electrospun from solvents such as urea,monochloroacetic acid, water, 2,2,2-trifluoroethanol, HFIP, orcombinations thereof. Elastin can be electrodeposited as a solution orsuspension in water, 2,2,2-trifluoroethanol, isopropanol, HFIP, orcombinations thereof, such as isopropanol and water. In one desirableembodiment, elastin is electrospun from a solution of 70% isopropanoland 30% water containing 250 mg/ml of elastin. Other lower orderalcohols, especially halogenated alcohols, may be used. Other solventsthat may be used or combined with other solvents in electrospinningnatural matrix materials include acetamide, N-methylformamide,N,N-dimethylformamide (DMF), dimethylsulfoxide (DMSO),dimethylacetamide, N-methyl pyrrolidone (NMP), acetic acid,trifluoroacetic acid, ethyl acetate, acetonitrile, trifluoroaceticanhydride, 1,1,1-trifluoroacetone, maleic acid, hexafluoroacetone.Organic solvents such as methanol, chloroform, and trifluoroethanol(TFE) and emulsifying agents.

The selection of a solvent is based in part on consideration ofsecondary forces that stabilize polymer-polymer interactions and thesolvent's ability to replace these with strong polymer-solventinteractions. In the case of polypeptides such as collagen, and in theabsence of covalent crosslinking, the principal secondary forces betweenchains are: (1) coulombic, resulting from attraction of fixed charges onthe backbone and dictated by the primary structure (e.g., lysine andarginine residues will be positively charged at physiological pH, whileaspartic or glutamic acid residues will be negatively charged); (2)dipole-dipole, resulting from interactions of permanent dipoles; thehydrogen bond, commonly found in polypeptides, is the strongest of suchinteractions; and (3) hydrophobic interactions, resulting fromassociation of non-polar regions of the polypeptide due to a lowtendency of non-polar species to interact favorably with polar watermolecules. Therefore, solvents or solvent combinations that canfavorably compete for these interactions can dissolve or dispersepolypeptides. For example, HFIP and TFE possess a highly polar OH bondadjacent to a very hydrophobic fluorinated region. While not wanting tobe bound by the following theories, it is believed that the alcoholportion can hydrogen bond with peptides, and can also solvate charges onthe backbone, thus reducing Coulombic interactions between molecules.Additionally, the hydrophobic portions of these solvents can interactwith hydrophobic domains in polypeptides, helping to resist the tendencyof the latter to aggregate via hydrophobic interactions. It is furtherbelieved that solvents such as HFIP and TFE, due to their lower overallpolarities compared to water, do not compete well for intramolecularhydrogen bonds that stabilize secondary structures such as an alphahelix. Consequently, alpha helices in these solvents are believed to bestabilized by virtue of stronger intramolecular hydrogen bonds. Thestabilization of polypeptide secondary structures in these solvents isbelieved desirable, especially in the cases of collagen and elastin, topreserve the proper formation of collagen fibrils duringelectrospinning.

In one embodiment, the solvent has a relatively high vapor pressure topromote the stabilization of an electrospinning jet to create a fiber asthe solvent evaporates. In embodiments involving higher boiling pointsolvents, it is often desirable to facilitate solvent evaporation bywarming the spinning or spraying solution, and optionally theelectrospinning stream itself, or by electrospinning in reducedatmospheric pressure. It is also believed that creation of a stable jetresulting in a fiber is facilitated by a high surface tension of thepolymer/solvent mixture.

Similar to conventional electrospinning, midair electrospinning can beused which employs the same experimental set-up as other electrospinningtechniques. However, in order to precipitate fibers before they reachthe grounded target, the distance from the needle to the grounded targetcan be increased. For example, increasing the distance from the 10-30 cmto a distance of 30-40 cm. The field strength can be maintained oraltered by increasing the applied potential at the needle tip.Increasing the distance from the needle tip to the grounded targetallows the polymer jet to experience a longer “flight time.” The addedflight time, allows the solvent to be completely evaporated from the jetallowing the fibers to fully develop.

By varying the composition of the fibers being electrospun, it will beappreciated that fibers having different physical or chemical propertiesmay be obtained. This can be accomplished either by spinning a liquidcontaining a plurality of components, each of which may contribute adesired characteristic to the finished product, or by simultaneouslyspinning fibers of different compositions from multiple liquid sources,that are then simultaneously deposited to form a matrix. The resultingmatrix comprises layers of intermingled fibers of different compounds.This plurality of layers of different materials can convey a desiredcharacteristic to the resulting composite matrix with each differentlayer providing a different property, for example one layer maycontribute to elasticity while another layer contributes to themechanical strength of the composite matrix. These methods can be usedto create tissues with multiple layers such as blood vessels.

The electrospun matrix has an ultrastructure with a three-dimensionalnetwork that supports cell growth, proliferation, differentiation anddevelopment. The spatial distance between the fibers plays an importantrole in cells being able to obtain nutrients for growth as well as forallowing cell-cell interactions to occur. Thus, in various embodimentsof the invention, the distance between the fibers may be about 50nanometers, about 100 nanometers, about 150 nanometers, about 200nanometers, about 250 nanometers, about 300 nanometers, about 350nanometers, about 600 nanometers, about 750 nanometers, about 800nanometers, about 850 nanometers, about 900 nanometers, about 950nanometers, about 1000 nanometers (1 micron), 10 microns, 10 microns, 50microns, about 100 microns, about 150 microns, about 200 microns, about250 microns, about 300 microns, about 350 microns, about 400 microns,about 450 microns, or about 500 microns. In various embodiments thedistance between the fibers may be less than 50 nanometers or greaterthan 500 microns and any length between the quoted ranges as well asintegers.

Additionally, in various embodiments of the invention, the fibers canhave a diameter of about 50 nanometers, about 100 nanometers, about 150nanometers, about 200 nanometers, about 250 nanometers, about 300nanometers, about 350 nanometers, about 600 nanometers, about 750nanometers, about 800 nanometers, about 850 nanometers, about 900nanometers, about 950 nanometers, about 1000 nanometers (1 micron), 50microns, about 100 microns, about 150 microns, about 200 microns, about250 microns, about 300 microns, about 350 microns, about 400 microns,about 450 microns, or about 500 microns, or the diameter may be lessthan 50 nanometers or greater than 500 microns and any diameter betweenthe quoted ranges as well as integers.

The pore size in an electrospun matrix can also be controlled throughmanipulation of the composition of the material and the parameters ofelectrospinning. In some embodiments, the electrospun matrix has a poresize that is small enough to be impermeable to one or more types ofcells. In one embodiment, the average pore diameter is about 500nanometers or less. In another embodiment, the average pore diameter isabout 1 micron or less. In another embodiment, the average pore diameteris about 2 microns or less. In another embodiment, the average porediameter is about 5 microns or less. In another embodiment, the averagepore diameter is about 8 microns or less. Some embodiments have poresizes that do not impede cell infiltration. In another embodiment, thematrix has a pore size between about 0.1 and about 100 μm². In anotherembodiment, the matrix has a pore size between about 0.1 and about 50μm². In another embodiment, the matrix has a pore size between about 1.0μm and about 25 μm. In another embodiment, the matrix has a pore sizebetween about 1.0 μm and about 5 μm. Infiltration can also beaccomplished with implants with smaller pore sizes. The pore size of anelectrospun matrix can be readily manipulated through control of processparameters, for example by controlling fiber deposition rate throughelectric field strength and mandrel motion, by varying solutionconcentration (and thus fiber size). Porosity can also be manipulated bymixing porogenic materials, such as salts or other extractable agents,the dissolution of which will leave holes of defined sizes in thematrix. The pore size can also be controlled by the amount ofcross-linking present in the matrix.

The mechanical properties of the matrix will depend on the polymermolecular weight and polymer type/mixture. It will also depend onorientation of the fibers (preferential orientation can be obtained bychanging speed of a rotating or translating surface during the fibercollection process), fiber diameter and entanglement. The cross-linkingof the polymer will also effect its mechanical strength after thefabrication process.

The electrospun matrix can be cross linked to increase its stability andstrength. The crosslinking can generally be conducted at roomtemperature and neutral pH conditions, however, the conditions may bevaried to optimize the particular application and crosslinking chemistryutilized. For crosslinking using the EDC chemistry with NHS in MES/EtOH,pH of from 4.0 to 8.0 and temperatures from 0° C. to room temperature(25° C.) for two hours, can be used. It is known that highertemperatures are unpreferred for this chemistry due to decomposition ofEDC. Similarly, basic pH (e.g., 8-14) is also unpreferred for thisreason when using this chemistry. Other crosslinking chemistries canalso be used for example, by soaking the electrospun matrix in 20%dextran solution (to reduce shrinking), followed by 1% glutaraldehydesolution. Yet other cross-linking chemistries involve usingpoly(ethylene glycol) (PEG) as a spacer in a crosslinking agent with anN-protected amino acid.

III. Synthetic Matrices

The invention also pertains to generating artificial tissue constructsby seeding cultured tissue cells onto or into available biocompatiblematrices. Biocompatible refers to materials that do not have toxic orinjurious effects on biological functions. Biodegradable refers tomaterial that can be absorbed or degraded in a patient's body.Representative materials for forming the biodegradable material includenatural or synthetic polymers, such as, collagen, poly(alpha esters)such as poly(lactate acid), poly(glycolic acid), polyorthoesters amdpolyanhydrides and their copolymers, which degraded by hydrolysis at acontrolled rate and are reabsorbed. These materials provide the maximumcontrol of degradability, manageability, size and configuration.Preferred biodegradable polymer materials include polyglycolic acid andpolyglactin, developed as absorbable synthetic suture material.

Polyglycolic acid and polyglactin fibers may be used as supplied by themanufacturer. Other biodegradable materials include, but are not limitedto, cellulose ether, cellulose, cellulosic ester, fluorinatedpolyethylene, phenolic, poly-4-methylpentene, polyacrylonitrile,polyamide, polyamideimide, polyacrylate, polybenzoxazole, polycarbonate,polycyanoarylether, polyester, polyestercarbonate, polyether,polyetheretherketone, polyetherimide, polyetherketone, polyethersulfone,polyethylene, polyfluoroolefin, polyimide, polyolefin, polyoxadiazole,polyphenylene oxide, polyphenylene sulfide, polypropylene, polystyrene,polysulfide, polysulfone, polytetrafluoroethylene, polythioether,polytriazole, polyurethane, polyvinyl, polyvinylidene fluoride,regenerated cellulose, silicone, urea-formaldehyde, or copolymers orphysical blends of these materials. The material may be impregnated withsuitable antimicrobial agents and may be colored by a color additive toimprove visibility and to aid in surgical procedures.

In some embodiments, attachment of the cells to the biocompatiblesubstrate is enhanced by coating the matrix with compounds such asbasement membrane components, agar, agarose, gelatin, gum arabic,collagens, fibronectin, laminin, glycosaminoglycans, mixtures thereof,and other materials having properties similar to biological matrixmolecules known to those skilled in the art of cell culture. Mechanicaland biochemical parameters ensure the matrix provide adequate supportfor the cells with subsequent growth and proliferation. Factors,including nutrients, growth factors, inducers of differentiation ordedifferentiation, products of secretion, immunomodulators, inhibitorsof inflammation, regression factors, biologically active compounds whichenhance or allow ingrowth of the lymphatic network or nerve fibers, anddrugs, can be incorporated into the matrix or provided in conjunctionwith the matrix. Similarly, polymers containing peptides such as theattachment peptide RGD (Arg-Gly-Asp) can be synthesized for use informing matrices.

Coating refers to coating or permeating a matrix with a material suchas, liquefied copolymers (poly-DL-lactide co-glycolide 50:50 80 mg/mlmethylene chloride) to alter its mechanical properties. Coating may beperformed in one layer, or multiple layers until the desired mechanicalproperties are achieved. These shaping techniques may be employed incombination, for example, a polymeric matrix can be weaved, compressionmolded and glued together. Furthermore different polymeric materialsshaped by different processes may be joined together to form a compositeshape. The composite shape can be a laminar structure. For example, apolymeric matrix may be attached to one or more polymeric matrixes toform a multilayer polymeric matrix structure. The attachment may beperformed by any suitable means such as gluing with a liquid polymer,stapling, suturing, or a combination of these methods. In addition, thepolymeric matrix may be formed as a solid block and shaped by laser orother standard machining techniques to its desired final form. Lasershaping refers to the process of removing materials using a laser.

The polymers can be characterized for mechanical properties such astensile strength using an Instron tester, for polymer molecular weightby gel permeation chromatography (GPC), glass transition temperature bydifferential scanning calorimetry (DSC) and bond structure by infrared(IR) spectroscopy; with respect to toxicology by initial screening testsinvolving Ames assays and in vitro teratogenicity assays, andimplantation studies in animals for immunogenicity, inflammation,release and degradation studies. In vitro cell attachment and viabilitycan be assessed using scanning electron microscopy, histology, andquantitative assessment with radioisotopes.

Substrates can be treated with additives or drugs prior to implantation(before or after the polymeric substrate is seeded with cells), e.g., topromote the formation of new tissue after implantation. Thus, forexample, growth factors, cytokines, extracellular matrix components, andother bioactive materials can be added to the substrate to promote grafthealing and formation of new tissue. Such additives will in general beselected according to the tissue or organ being reconstructed oraugmented, to ensure that appropriate new tissue is formed in theengrafted organ or tissue (for examples of such additives for use inpromoting bone healing, see, e.g., Kirker-Head, C. A. Vet. Surg. 24 (5):408-19 (1995)). For example, vascular endothelial growth factor (VEGF,see, e.g., U.S. Pat. No. 5,654,273 herein incorporated by reference) canbe employed to promote the formation of new vascular tissue. Growthfactors and other additives (e.g., epidermal growth factor (EGF),heparin-binding epidermal-like growth factor (HBGF), fibroblast growthfactor (FGF), cytokines, genes, proteins, and the like) can be added inamounts in excess of any amount of such growth factors (if any) whichmay be produced by the cells seeded on the substrate. Such additives arepreferably provided in an amount sufficient to promote the formation ofnew tissue of a type appropriate to the tissue or organ, which is to berepaired or augmented (e.g., by causing or accelerating infiltration ofhost cells into the graft). Other useful additives include antibacterialagents such as antibiotics.

The biocompatible substrate may be shaped using methods such as, solventcasting, compression molding, filament drawing, meshing, leaching,weaving and coating. In solvent casting, a solution of one or morepolymers in an appropriate solvent, such as methylene chloride, is castas a branching pattern relief structure. After solvent evaporation, athin film is obtained. In compression molding, the substrate is pressedat pressures up to 30,000 pounds per square inch into an appropriatepattern. Filament drawing involves drawing from the molten polymer andmeshing involves forming a mesh by compressing fibers into a felt-likematerial. In leaching, a solution containing two materials is spreadinto a shape close to the final form of the tissue. Next a solvent isused to dissolve away one of the components, resulting in poreformation. (See Mikos, U.S. Pat. No. 5,514,378, hereby incorporated byreference).

In nucleation, thin films in the shape of the tissue are exposed toradioactive fission products that create tracks of radiation damagedmaterial. Next, the polycarbonate sheets are etched with acid or base,turning the tracks of radiation-damaged material into pores. Finally, alaser may be used to shape and burn individual holes through manymaterials to form a tissue structure with uniform pore sizes. Thesubstrate can be fabricated to have a controlled pore structure thatallows nutrients from the culture medium to reach the deposited cellpopulation. In vitro cell attachment and cell viability can be assessedusing scanning electron microscopy, histology and quantitativeassessment with radioisotopes.

Thus, the substrate can be shaped into any number of desirableconfigurations to satisfy any number of overall system, geometry orspace restrictions. The matrix can be shaped to different sizes toconform to the necessary structures of different sized patients.

A substrate can also be permeated with a material, for example liquifiedcopolymers (poly-DL-lactide co-glycolide 50:50 80 mg/ml methylenechloride) to alter its mechanical properties. This can be performed bycoating one layer, or multiple layers until the desired mechanicalproperties are achieved.

The substrate can also be treated or seeded with various factors andproteins to control the degradation/absorption of the matrix in thesubject. For instance, if the cells seeded within the substrate areslow-growing, then it is beneficial to maintain the matrix integrity fora long enough period of time to allow the cells enough time toregenerate and grow. On the other hand, if the cells are able to quicklyreproduce and grow, then a short lived substrate could be desirable.Varying the concentration of aprotinin additives, aminocaproic acid,tranxemic acid, or similar fibrinolytic inhibitors or the degree ofchemical cross-linking in the matrix could be used to precisely controlthis variable. The substrate could also be seeded with varying growthfactors such as angiogenesis factor to promote a growth of blood vesselsupon implantation.

All types of matrices are preferably shaped into a tubular orcylindrical structures with length and diameter dimensions selected tocorrespond with digits or limbs of the subject. All types of matricescan be joined or bonded together or to a structural bone of the subjectusing standard techniques such as suturing, and gluing with biologicalglue. Suturing can involve known techniques using absorbable syntheticsuture material such as the biocompatible polymer is polyglactin andpolyglycolic acid, manufactured as Vicryl™ by Ethicon Co., Somerville,N.J. (See e.g., Craig P. H., Williams J. A., Davis K. W., et al.: ABiological Comparison of Polyglactin 910 and Polyglycolic Acid SyntheticAbsorbable Sutures. Surg. 141; 1010, (1975)). Other methods of joininginvolve using biological glues. Biological glues can adhere to tissues,attach them to each other, or attach them to other structures on thebody in a few minutes, without using staples or sutures. These glues areeliminated, in general after the cicatrization of the wound, bybiodegradation, resorption or by simple detachment in the form of scabs.

Various technologies have been developed for the formulation of tissueadhesives. Some of them are of synthetic origin, such as the glues basedon cyanoacrylates (2-butyl cyanoacrylate, 2-octyl cyanoacrylate), or onsynthetic polymers, and others contain biological materials such ascollagen or fibrin (See e.g., U.S. Pat. No. 5,844,016, U.S. Pat. No.5,874,500; U.S. Pat. No. 5,744,545; U.S. Pat. No. 5,550,187 and U.S.Pat. No. 6,730,299).

IV. Intelligent Scaffolds

The invention pertains to using “intelligent” scaffolds that recruitcells, and encourage the formation of a vascular support and innervationvia release of chemical factors. These intelligent scaffolds possess thenecessary ultrastructural, biomechanical and biological characteristicsare required for cell attachment, survival, neovascularization,innervation, and tissue maturation even in an environment with a limitedblood supply. The intelligent scaffold can be created by adding anysuitable therapeutic or biological agent such as genetic material,growth factors, cytokines, enzymes can be used to create an intelligentscaffold. The therapeutic or biological agent can recruit cells andfactors that encourage growth and proliferation of cells seeded on thescaffold. These intelligent scaffolds can also be used to encouragenerve growth and innervation.

Examples of a therapeutic or biological agent include, but are notlimited to proteins growth factors, antibodies, nucleic acids molecules,carbohydrates, and the like. Growth factors useful in the presentinvention include, but are not limited to, vascular endothelial growthfactor (VEGF), nerve growth factors (NGF) including NGF 2.5 s, NGF 7.0 sand beta NGF and neurotrophis, transforming growth factor-alpha (TGF-α),transforming growth factor-beta (TGF-β), platelet-derived growth factors(PDGF), fibroblast growth factors (FGF), including FGF acidic isoforms 1and 2, FGF basic form 2 and FGF 4, 8, 9 and 10, brain derivedneurotrophic factor, cartilage derived factor, bone growth factors(BGF), basic fibroblast growth factor, insulin-like growth factor (IGF),granulocyte colony stimulating factor (G-CSF), insulin like growthfactor (IGF) I and II, hepatocyte growth factor, glial neurotrophicgrowth factor (GDNF), stem cell factor (SCF), keratinocyte growth factor(KGF), transforming growth factors (TGF), including TGFs alpha, beta,beta1, beta2, beta3, skeletal growth factor, bone matrix derived growthfactors, and bone derived growth factors and mixtures thereof.

Cytokines useful in the present invention include, but are not limitedto, cardiotrophin, stromal cell derived factor, macrophage derivedchemokine (MDC), melanoma growth stimulatory activity (MGSA), macrophageinflammatory proteins 1 alpha (MIP-1 alpha), 2, 3 alpha, 3 beta, 4 and5, IL-1, IL-2, IL-3, IL-4, IL-5, IL-6, IL-7, IL-8, IL-9, IL-10, IL-11,IL-12, IL-13, TNF-alpha, and TNF-beta. Immunoglobulins useful in thepresent invention include, but are not limited to, IgG, IgA, IgM, IgD,IgE, and mixtures thereof.

Other molecules useful as therapeutic or biological agents include, butare not limited to, growth hormones, leptin, leukemia inhibitory factor(LIF), endostatin, thrombospondin, osteogenic protein-1, bonemorphogenetic proteins 2 and 7, osteonectin, somatomedin-like peptide,osteocalcin, interferon alpha, interferon alpha A, interferon beta,interferon gamma, interferon I alpha.

Embodiments involving amino acids, peptides, polypeptides, and proteinsmay include any type or combinations of such molecules of any size andcomplexity. Examples include, but are not limited to structuralproteins, enzymes, and peptide hormones. These compounds can serve avariety of functions. In some embodiments, the matrix may containpeptides containing a sequence that suppresses enzyme activity throughcompetition for the active site. In other applications antigenic agentsthat promote an immune response and invoke immunity can be incorporatedinto a construct. In substances such as nucleic acids, any nucleic acidcan be present. Examples include, but are not limited todeoxyribonucleic acid (DNA), and ribonucleic acid (RNA). Embodimentsinvolving DNA include, but are not limited to, cDNA sequences, naturalDNA sequences from any source, and sense or anti-sense oligonucleotides.For example, DNA can be naked (e.g., U.S. Pat. Nos. 5,580,859;5,910,488) or complexed or encapsulated (e.g., U.S. Pat. Nos. 5,908,777;5,787,567). DNA can be present in vectors of any kind, for example in aviral or plasmid vector. In some embodiments, nucleic acids used willserve to promote or to inhibit the expression of genes in cells insideand/or outside the electrospun matrix. The nucleic acids can be in anyform that is effective to enhance its uptake into cells.

Factors involved in bone formation include hormones such as estrogen,calcitonin, and parathyroid hormone (PTH); growth factors such as bonemorphologenic protein (BMP); and chemicals such as active vitamin D,calcium preparations, and vitamin K2. Estrogen, calcitonin, activevitamin D, and calcium preparations are used as medicine for controllingbone mass in osteoporosis or similar cases.

The intelligent scaffolds can be created by functionalizing the scaffoldto incorporate a contrast enhancing agent (e.g., gadolinium) or ananoparticles such as a quantum dot coupled to a biological ortherapeutic agent. Quantum dots are a semiconductor nanocrystal withsize-dependent optical and electronic properties. In particular, theband gap energy of a quantum dot varies with the diameter of thecrystal. The average diameter of the QDs may be between about 1 to about100 nm, between about 10-80 nm, and between about 25-40 nm. The coupledagent can be released by application of energy such as near infrared(NIR) irradiation from a laser source, which causes the bonds betweenthe agent and the QD to break and thus releases the agent. This allowsthe release of the agent to be controlled by triggering its release uponapplication of energy. Quantum dots have been used as photostablebiological fluorescent tags, semiconductors, and thermal therapy. Thehigh transmission, scattering-limited attenuation, and minimal heatingeffects of quantum dots makes these suitable for the coupling oftherapeutic/biological agents. In one embodiment, NIR CdSe quantum dots(Evident Technologies) can be used. These QDs have an optical absorptionrange of 700-1000 nm. NIR energy within this spectral region has beenshown to penetrate tissue at depths up to 23 cm with no observabledamage to the intervening tissue.

A matrix functionalized with a QD coupled to a therapeutic or biologicalagent can be used for controlled release of the therapeutic orbiological agent at a target in the subject. The therapeutic orbiological agent can be released by application of energy at a desiredwavelength such as near infrared irradiation. Due to localized heatingof the QD, ultrastructural changes cause the release of the coupledagent. The release kinetics can be varied according to the type of QDused and the wavelength of irradiation. The release kinetics can also bevaried by altering the intensity and time of irradiation. For example, aQD (e.g., CdSe QD from Evident Technologies) coupled to encapsulatedheparin can be incorporated into an electrospun matrix. Upon applicationof near infrared radiation at a wavelength of 700-1000 nm, the heparinis released in a controlled manner, as described in the examples below.

The studies in the examples section demonstrate the burst release ofheparin over time when quantum dot conjugated heparin nanoparticles wereirradiated by NIR irradiation. This system allows medical personnel totune therapeutic/biological agent release rates post-operatively.

The emission spectra of quantum dots have linewidths as narrow as 25-30nm depending on the size heterogeneity of the sample, and lineshapesthat are symmetric, gaussian or nearly gaussian with an absence of atailing region. The combination of tunability, narrow linewidths, andsymmetric emission spectra without a tailing region provides for highresolution of multiply-sized quantum dots within a system and allowssimultaneous examination of a variety of biological moieties tagged withQDs.

In addition, the range of excitation wavelengths of the quantum dots isbroad and can be higher in energy than the emission wavelengths of allavailable quantum dots. Consequently, this allows the simultaneousexcitation of all quantum dots in a system with a single light source.The ability to control the size of QDs enables one to construct QDs withfluorescent emissions at any wavelength in the UV-visible-IR region.Therefore, the colors (emissions) of QDs are tunable to any desiredspectral wavelength. Furthermore, the emission spectra of monodisperseQDs have linewidths as narrow as 25-30 nm. The linewidths are dependenton the size heterogeneity of QDs in each preparation. In one embodiment,the QDs emit light in the ultraviolet wavelengths. In anotherembodiment, the QDs emit light in the visible wavelengths. In otherembodiments, the QDs emit light in the near-infrared and the infraredwavelengths. Color of the light emitted by the QDs may be size-tunableand excitation energy tunable.

Many QDs are constructed of elements from groups II-VI, III-V and IV ofthe periodic table. They exhibit quantum confinement effects in theirphysical properties, and can be used in the composition of theinvention. Exemplary materials suitable for use as quantum dots include,but are not limited to, ZnS, ZnSe, ZnTe, CdS, CdSe, CdTe, GaN, GaP,GaAs, GaSb, InP, InAs, InSb, AlS, AlP, AlAs, AlSb, PbS, PbSe, Ge, and Siand ternary and quaternary mixtures thereof. The quantum dots mayfurther include an overcoating layer of a semiconductor having a greaterband gap.

In particular, the therapeutic or biological agent and the nanoparticles(e.g., quantum dot) can be entrapped or encapsulated to produce“nanocapsules.” These nanocapsules containing the agent and thenanoparticle can be produce standard encapsulating techniques.Microencapsulation of agents generally involves three steps: (a)generating microcapsules enclosing the agents (e.g., by forming dropletsof cell-containing liquid alginate followed by exposure to a solution ofcalcium chloride to form a solid gel), (b) coating the resulting gelledspheres with additional outer coatings (e.g., outer coatings comprisingpolylysine and/or polyomithine) to form a semipermeable membrane; and(c) liquefying the original core gel (e.g., by chelation using asolution of sodium citrate). The three steps are typically separated bywashings in normal saline.

Alginates are linear polymers of mannuronic and guluronic acid residues.Monovalent cation alginate salts, e.g., Na-alginate, are generallysoluble. Divalent cations such as Ca²⁺, Ba²⁺ or Sr²⁺ tend to interactwith guluronate, providing crosslinking and forming stable alginategels. Alginate encapsulation techniques typically take advantage of thegelling of alginate in the presence of divalent cation solutions.Alginate encapsulation of agent-nanoparticles generally involvessuspending the agent-nanoparticles to be encapsulated in a solution of amonovalent cation alginate salt, generating droplets of this solution,and contacting the droplets with a solution of divalent cations. Thedivalent cations interact with the alginate at the phase transitionbetween the droplet and the divalent cation solution, resulting in theformation of a stable alginate gel matrix being formed. A variation ofthis technique is reported in U.S. Pat. No. 5,738,876, where the cell issuffused with a solution of multivalent ions (e.g., divalent cations)and then suspended in a solution of gelling polymer (e.g., alginate), toprovide a coating of the polymer.

Another method of microencapsulating agent-nanoparticles is thealginate-polyamino acid technique. Cells are suspended in sodiumalginate in saline, and droplets containing islets are produced.Droplets of cell-containing alginate flow into calcium chloride insaline. The negatively charged alginate droplets bind calcium and form acalcium alginate gel. The microcapsules are washed in saline andincubated with poly-L-lysine (PLL) or poly-L-ornithine (or combinationsthereof); the positively charged poly-l-lysine and/or poly-L-ornithinedisplaces calcium ions and binds (ionic) negatively charged alginate,producing an outer poly-electrolyte membrane. A final coating of sodiumalginate may be added by washing the microcapsules with a solution ofsodium alginate, which ionically bonds to the poly-L-lysine and/orpoly-L-ornithine layer. See U.S. Pat. No. 4,391,909 to Lim et al (allU.S. patents referenced herein are intended to be incorporated herein intheir entirety). This technique produces what has been termed a“single-wall” microcapsule. Preferred microcapsules are essentiallyround, small, and uniform in size. (Wolters et al., J. Appli Biomater.3:281 (1992)).

The alginate-polylysine microcapsules can also then be incubated insodium citrate to solubilize any calcium alginate that has not reactedwith poly-l-lysine, i.e., to solubilize the internal core of sodiumalginate containing the islet cells, thus producing a microcapsule witha liquefied cell-containing core portion. See Lim and Sun, Science210:908 (1980). Such microcapsules are referred to herein as having“chelated”, “hollow” or “liquid” cores. A “double-wall” microcapsule isproduced by following the same procedure as for single-wallmicrocapsules, but prior to any incubation with sodium citrate, themicrocapsules are again incubated with poly-l-lysine and sodiumalginate.

Many alternative techniques used for encapsulating agents are known inthe art and can be used with this invention. U.S. Pat. No. 5,084,350discloses microcapsules enclosed in a larger matrix, where themicrocapsules are liquefied once the microcapsules are within the largermatrix. Tsang et al., U.S. Pat. No. 4,663,286 discloses encapsulationusing an alginate polymer, where the gel layer is cross-linked with apolycationic polymer such as polylysine, and a second layer formed usinga second polycationic polymer (such as polyornithine); the second layercan then be coated by alginate. U.S. Pat. No. 5,762,959 to Soon-Shionget al. discloses a microcapsule having a solid (non-chelated) alginategel core of a defined ratio of calcium/barium alginates, with polymermaterial in the core. U.S. Pat. Nos. 5,801,033 and 5,573,934 to Hubbellet al. describe alginate/polylysine microspheres having a finalpolymeric coating (e.g., polyethylene glycol (PEG)); Sawhney et al.,Biomaterials 13:863 (1991) describe alginate/polylysine microcapsulesincorporating a graft copolymer of poly-l-lysine and polyethylene oxideon the microcapsule surface, to improve biocompatibility; U.S. Pat.No.5,380,536 describes microcapsules with an outermost layer of watersoluble non-ionic polymers such as polyethylene(oxide). U.S. Pat. No.5,227,298 to Weber et al. describes a method for providing a secondalginate gel coating to cells already coated with polylysine alginate;both alginate coatings are stabilized with polylysine. U.S. Pat. No.5,578,314 to Weber et al. provides a method for microencapsulation usingmultiple coatings of purified alginate. U.S. Pat. No. 5,693,514 toDorian et al. reports the use of a non-fibrogenic alginate, where theouter surface of the alginate coating is reacted with alkaline earthmetal cations comprising calcium ions and/or magnesium ions, to form analkaline earth metal alginate coating. The outer surface of the alginatecoating is not reacted with polylysine. U.S. Pat. No. 5,846,530 toSoon-Shiong describes microcapsules containing cells that have beenindividually coated with polymerizable alginate, or polymerizablepolycations such as polylysine, prior to encapsulation.

In one embodiment, heparin is coupled to the nanoparticle and thecontrolled release kinetics of heparin can be monitored. One skilled inthe art will appreciate that the control release kinetics depend on thecapsulation parameters including nanocapsule size, heparin and quantumdot loading, and polymer composition. The mean diameter of thenanocapsules depends on the mixing velocity of the preparation processand viscosity of the preparation media. Nanocapsule size can be reducedby exposing the preparation to sonication over a range of about 30second to about 120 seconds, increasing the sonication intensity fromabout 5 watts to about 20 watts, or by varying the ratios of organicpolymer phase to aqueous heparin phase. Nanocapsule sizes can becharacterized by scanning electron microscopy (SEM), coulter counter,and light scattering.

In one embodiment, the heparin can be conjugated to quantum dots byusing an EDC/NHS chemical method. Various concentrations of heparin(ranging form 10-30 weight % polymer) and quantum dots canbe used todetermine optimal loading efficiency.

For polymer encapsulation, FDA approved biodegradable polymers (PLA,PLGA, PCL) can be used for the control of encapsulation and degradationof the nanocapsules in vivo.

The examples show that a burst of heparin release occurs using abroadband infrared (IR) source. Using measured quantities of QD-Heparinnanocapsules (NC) suspended in a physiological buffer, the influence ofvarying wavelengths, intensities, and irradiation times on the releasekinetics can be determined. In one embodiment, the wavelength ofirradiation used on the QD-Heparin can be in the near-infraredwavelength range, such as 700 nm, 800 nm, and 900 nm, using a filteredxenon source. The intensity of irradiation energy can be adjusted inincremental steps from 0 (control), 1 mW/cm², 10 mW/cm², 100 mW/cm², 1W/cm², and 10 W/cm². The irradiation time can also be varied todetermine the optimal irradiation time at each effective powerintensity. The irradiation time can vary from 0 (control), 10, 60, 300,and 600 seconds of exposure.

The encapsulated QD-heparin will be released upon near infra-red (NIR)irradiation due to localized heating of the quantum dots which inducesultrastructural changes in the nanocapsules. The release kinetics willbe varied at the target site by modulating the intensity and time of NIRirradiation to produce a controlled release of heparin. The quantitativemeasurement of heparin released from the nanocapsules can be measuredover time (2, 4, 6, 12, and 24 hours and daily thereafter up to 30 days)and measured for its anti-factor Xa activity with a syntheticchromogenic substrate using a kit Rotachrom (Diagnostica Stago Inc).

The scaffold may also be functionalized with an image enhancing agent tomonitor growth and assimilation of the construct in vivo. In anotheraspect, the invention pertains to monitoring tissue remodeling a tissueengineered construct. Remodeling that takes place too slowly can resultin pathologic response of surrounding tissues and compliance mismatch ofthe vessel. Rapid remodeling can result in premature failure of theengineered construct. Magnetic Resonance Imaging (MRI) is a powerful,non-invasive technique that can be used long term for monitoring theremodeling process. Nanoparticles (e.g., QD, image enhancing agents) canbe easily bound to matrices, and also embedded within nanofibers ofelectrospun matrices. The nanoparticles provide high MRI contrast, anddue to their small size, will not interfere with normal biologicalprocesses. Organolanthanide complexes containing paramagnetic metalssuch as gadolinium (Gd) have been known to cause distortion in anelectromagnetic field. When the protons in water interact with thisdistorted field, their magnetic properties significantly change suchthat they can be detected by MRI. The Examples demonstrate the enhancedimaging observed using MRI contrast with Gd functionalized nanoparticlesbound to the surface and/or incorporated into the vascular matrices ornanocapsules. Other examples of contrast enhancing agents include, butare not limited to, rare earth metals such as, cerium, samarium,terbium, erbium, lutetium, scandium, barium, bismuth, cerium,dysprosium, europium, hafnium, indium, lanthanum, neodymium, niobium,praseodymium, strontium, tantalum, ytterbium, yttrium, and zirconium.

In one embodiment, the agents are joined to the matrix by peptide bonds.For example, nanoparticles can be incorporated as part of the matrixusing EDC (1-ethyl-3(3-dimethly aminopropyl)carbodiimide) and sulfo-NHS(N-hydrocyl-sulfo-succinimide) to form peptide bonds. Various other knowtechniques can be used as described, for example, in Heumanson,Bioconjugate Techniques, Academic Press San Diego, Calif., 1996, hereinincorporated by reference. For external functionalization, a peptidebond can be created between the matrix and carboxylated gadoliniumnanoparticles using the EDC/sulpho-NHS method to form peptide bondsbetween the carboxylates and amino groups. The quantum dot coupled to atherapeutic/biological agent, a contrast enhancing agent, e.g.,gadolinium, or both, can also be added internally to an electrospunmatrix by incorporating each component into the solution with at leastone natural compound and at least one synthetic compound. For example,solutions containing collagen I, elastin and PLGA, successfullyincorporated the contrast enhancing agent gadolinium uponelectrospinning as described in the Examples. The incorporation of thegadolinium into the matrix can be observed in vitro and in vivo usingdetection methods such as magnetic resonance imaging (MRI). Thus, amatrix functionalized with a contrasting agent allows the degradation ofthe matrix to be monitored.

Any type of functionalization method can be used. Examples of somepossible functionalization chemistries include, but are not limited to,esterification (e.g., with acyl halides, acid anhydrides, carboxylicacids, or esters via interchange reactions), ether formation (forexample, via the Williamson ether synthesis), urethane formation viareactions with isocyanates, sulfonation with, for example,chlorosulfonic acid, and reaction of b-sulfato-ethylsulfonyl aniline toafford an amine derivative that can be converted to a diazo for reactionwith a wide variety of compounds. Such chemistries can be used to attacha wide variety of substances to the electrospun matrix, including butnot limited to crown ethers (Kimura et al., (1983) J. Polym. Sci. 21,2777), enzymes (Chase et al. (1998) Biotechnol. Appl. Biochem., 27,205), and nucleotides (Overberger et al. (1989) J. Polym. Sci. 27,3589).

V. Composite Scaffolding

Numerous materials can be used as scaffolds for bone tissuereconstruction as described above. These include metals, ceramics andpolymers from biologic and synthetic origins. Synthetic materials, suchas hydroxyapatite (HA), tricalcium phosphate (TCP), calcium sulfate,poly(lactide-co-glycolide) (PLGA), polyglycolide (PGA) and polylactide(PLA) can be used either alone or in combination with naturally derivedmaterials including collagen, chitosan, starch and silk fibroin. Thesematerials are designed to serve as a bone substitute or as anenhancement for the bone healing process. Among the several commerciallyavailable bone graft materials, Collagraft (collagen-hydroxyapatitecomposite scaffold) is currently the most commonly used materialclinically.

In some embodiments, a cell-based approach can be used in bone tissueregeneration. While many biomaterials serve as a scaffold that augmentsthe body's ability to heal itself, a tissue engineering approach usescells added to a scaffold to achieve formation of bone tissue. Existingbiomaterials, such as Collagraft, may not be ideal for use with cellsdue to their physical and structural configuration, which includes lowcell adhesion, poor cell infiltration and brittleness. PLGA has beenused as a scaffold for bone tissue engineering due to its favorablephysical properties. However, the surface chemistry of PGLA does notfully promote cell adhesion and proliferation due to its hydrophobicnature.

In one aspect, the invention discloses composite bone scaffoldscomprising a biodegradable synthetic polymer and a naturally derivedcollagen matrix. Examples of synthetic polymers include, but are notlimited to, poly(lactide-co-glycolide) (PLGA), poly(lactide) (PLA),poly(glycolic acid) (PGA), poly(caprolactone), polycarbonates,polyamides, polyanhydrides, polyamino acids, polyortho esters,polyacetals, polycyanoacrylates, degradable polyurethanes,hydroxyapatite (HA), tricalcium phosphate (TCP), and calcium sulfate.Examples of a naturally derived collagen matrix suitable for the presentinvention include, but are not limited to, submucosa, such as bladder,stomach, esophagus, stomach, small intestine, large intestine (colon)submucosa. The submucosa is a layer of interstitial protein thatsupports blood vessels, which supply the mucosa with nutrients and thelymph nodes which aid in the removal of waste products. The submucosaserves an important function, and is produced as the interface betweenthe mucosa and the detrusor. Many organs are made up of multiple layersof different tissues. Depending on the functional role of the organ,different tissues confer different properties to the organ. For example,the bladder has three main layers of tissue: the mucosa, submucosa anddetrusor. The submucosa can be harvested from a mammal (i.e., human,cow, pig, etc.) as described below.

The biocompatible composite scaffolding system preferably has asubstantially uniform porous structures having an average pore diameterranging from about 50 to about 250 μm, preferably from about 90 to about150 μm, and most preferably from about 110 to about 130 μm. The porosityof the biocompatible composite scaffolding is greater than about 50%,preferably greater than about 80%, more preferably greater than about90%, and most preferably greater than about 95%.

In another aspect, the invention provides a method of fabricating abiocompatible composite scaffolding system capable of providingstructural support for regenerated bone tissue comprising selecting apore size, obtaining porogens of said selected size, adding the porogensto a solution containing a synthetic polymer and submucosa and mixing todistribute said porogens and form a composite scaffold, drying saidcomposite scaffold to remove residual solvent, and removing saidporogens from said composite scaffold. The method can further includesieving porogens through a sieve to obtain said selected size porogens.The step of removing said porogens from said composite scaffold cancomprise immersion in water. As used here, the term “porogen” refers toany soluble particulate. Some exemplary porogen materials suitable foruse in the present invention are selected from the group consisting ofsalts including, but not limited to sodium chloride, potassium chloride,sodium fluoride, potassium fluoride, sodium iodide, sodium nitrate,sodium sulfate, sodium iodate, and mixtures thereof, other water solublechemicals such as sodium hydroxide, sugars including, but not limited tosaccharin, glucose, fructose, other water soluble sugars, and mixturesthereof, waxes paraffin, beeswax, other waxes, and mixtures thereof,gelatins, naphthalene, natural or synthetic water soluble polymers,natural or synthetic non-water soluble polymers, degradable polymers,non-degradable polymers, partially degradable polymers, and mixturesthereof. The porogen materials can be formed into any shape as desiredand/or necessary. However, in the preferred embodiment, thepredetermined shape is selected from the group consisting of cubic orother geometrically shaped crystals, spheres, fibers, discs, regulargeometric shapes, irregular geometric shapes, and mixtures thereof. Apreferred porogen is sodium chloride. The synthetic polymer can beselected from the group comprising poly(lactide-co-glycolide) (PLGA),poly(lactide) (PLA), poly(glycolic acid) (PGA), poly(caprolactone),polycarbonates, polyamides, polyanhydrides, polyamino acids, polyorthoesters, polyacetals, polycyanoacrylates, degradable polyurethanes,hydroxyapatite (HA), tricalcium phosphate (TCP), and calcium sulfate. Ina preferred embodiment, the synthetic polymer ispoly(D,L-lactide-co-glycolide) (PLGA) and the submucosa is bladdersubmucosa (BSM). The preferred weight ratio of PLGA and BSM to NaCl isabout 1:10, but can range from about 1:5 to about 1:30.

The composite scaffolds are configured to accommodate cells and designedto provide adequate structural support. The biological activity,physical and structural properties of the scaffolds for their use inbone tissue regeneration using two different cell types were confirmedas shown in Example 28. Primary mature osteoblasts and stem cells formbone tissues when implanted in vivo or differentiated into bone cells.These cell types were used to determine their ability to survive, adhereand proliferate on the composite scaffolds. Cell accommodation, adhesionand proliferation were approximately 80% higher in the BSM-PLGAcomposite scaffolds, as compared to the control scaffolds using thesetwo cell types. Fabrication of the scaffolds with an appropriate poresize, porosity and surface hydrophilicity resulted in abundant cellaccommodation with increased cell proliferation. The incorporation ofBSM significantly improved the biological activities of the scaffold,while maintaining physical and structural stability.

Creation of bone tissue using cells requires a scaffold that serves as acell carrier which would provide structural support until bone tissueforms in vivo. The scaffold for bone tissue engineering should bebiocompatible and possess mechanical stability, a controlled degradationrate, hydrophilic surface chemistry and an appropriate porosity for cellaccommodation. In one embodiment of this invention, a composite scaffoldfor bone regeneration is disclosed that meets these criteria byhybridizing BSM as a natural bioactive material with synthetic PLGApolymers. Example 28 demonstrates that the biodegradable syntheticpolymer and naturally derived collagen matrix composite scaffolds arenon-toxic, easily fabricated, and provide structural support withabundant pores with good interconnectivity.

A scaffold material for bone regeneration should be biocompatible andsafe for implantation. Cellular interactions of the BSM-PLGA compositescaffold were tested by using the two widely accepted complementaryassays; cell viability and mitochondrial metabolic activity. The cellviability assay using Neutral Red, which is based on dye incorporationinto lysosomes, and the MTT assay, which is based on the intact activityof a mitochondrial enzyme, demonstrated that the BSM-PLGA compositescaffolds are safe.

VI. Culturing Cells

To engineer a limb tissue composed of a patient's own cells, a reliablecell expansion system for each cell type was established. Stem andprogenitor cells in particular, muscle progenitor cells (MPC) are anideal cell source for tissue engineering. These cells have the capacityof self renewal and are multipotent i.e. they are able to differentiateinto a wide variety of cell types and tissues. Adult stem and progenitorcells can be isolated from various sources. In addition to “tissuespecific” stem cells found to reside within a number of adult tissues,the bone marrow stroma provides a “universal” source of stem cells thatparticipate in regenerative processes of many tissues. Stem andprogenitor cells from various stages of development, have been usedincluding embryonic, fetal and adult, from multiple tissue sources. Asystem has been established to differentiate adult stem cells into bone,muscle, fat, endothelial and nerve cells. These cells were furthertransplanted and formed functional tissues in vivo.

Bone marrow stem cells can also be used for the engineering of bonetissues of the digit. Bone marrow cells can be easily obtained through asimple needle aspiration and density gradient centrifugation. The bonemarrow cells will be expanded in culture, characterized and induced tobecome bone cells in the presence of osteogenic supplements. Thephenotypic and functional characteristics of the bone cells can bedetermined by evaluation of alkaline phosphatase production andactivity, calcium deposition and mineralization.

Muscle cells e.g. myoblasts, can be obtained from a skeletal muscletissue biopsy. MPC migrated from the single fibers will be grown andexpanded. Phenotypical and functional assessment of the myoblastsincludes testing their ability to form myotubes, as well as Westernblot, RT-PCR and immunohistochemical analyses. Myoblasts can be seededon muscle scaffolds for tissue formation in vitro. The engineered boneand muscle composite tissues can be assessed for their structural,physical and physiological characteristics.

There are five main types of bone cells in bone tissue. Osteogenic cellsrespond to traumas, such as fractures, by giving rise to bone-formingcells and bone-destroying cells. Osteoblasts (bone-forming cells)synthesize and secrete unmineralized ground substance and are found inareas of high metabolism within the bone. Osteocytes are mature bonecells made from osteoblasts that have made bone tissue aroundthemselves. These cells maintain healthy bone tissue by secretingenzymes and controlling the bone mineral content; they also control thecalcium release from the bone tissue to the blood. Osteoclasts are largecells that break down bone tissue. They are very important to bonegrowth, healing, and remodeling. The last type of cells are bone-liningcells. These are made from osteoblasts along the surface of most bonesin an adult. Bone-lining cells are thought to regulate the movement ofcalcium and phosphate into and out of the bone. For cartilage cells,chondrocytes and chondroblasts can be used.

The artificial tissue can be created by using allogenic cell populationsderived from the subject's own tissue. The artificial tissue can also bexenogenic, where cell populations are derived from a mammalian speciesthat are different from the subject. For example, muscle,cartilage-forming or bone-forming tissue cells can be derived frommammals such as monkeys, dogs, cats, mice, rats, cows, horses, pigs,goats and sheep.

The isolated cells are preferably cells obtained by a swab or biopsy,from the subject's own tissue. A biopsy can be obtained by using abiopsy needle under a local anesthetic, which makes the procedure quickand simple. The small biopsy core of the isolated tissue can then beexpanded and cultured to obtain the tissue cells. Cells from relativesor other donors of the same species can also be used with appropriateimmunosuppression.

Methods for the isolation and culture of cells are discussed byFreshney, Culture of Animal Cells. A Manual of Basic Technique, 2d Ed.,A. R. Liss, Inc., New York, 1987, Ch. 9, pp. 107-126. Cells may beisolated using techniques known to those skilled in the art. Forexample, the tissue can be cut into pieces, disaggregated mechanicallyand/or treated with digestive enzymes and/or chelating agents thatweaken the connections between neighboring cells making it possible todisperse the tissue into a suspension of individual cells withoutappreciable cell breakage. If necessary, enzymatic dissociation can beaccomplished by mincing the tissue and treating the minced tissue withany of a number of digestive enzymes either alone or in combination.These include but are not limited to trypsin, chymotrypsin, collagenase,elastase, and/or hyaluronidase, DNase, pronase, and dispase. Mechanicaldisruption can also be accomplished by a number of methods including,but not limited to, scraping the surface of the tissue, the use ofgrinders, blenders, sieves, homogenizers, pressure cells, or insonatorsto name but a few.

Cell types include, but are not limited to, progenitor cells isolatedfrom the peripheral blood or bone marrow cells that can be induced todifferentiate into different cells such as muscle cells, bone cells orcartilage cells, stem cells, committed stem cells, and/or differentiatedcells may be used. Also, depending on the type of tissue or organ beingmade, specific types of committed stem cells can be used. For instance,myoblast cells can be used to build various muscle structures. Othercells include, but are not limited to, endothelial cells, muscle cells,smooth muscle cells, fibroblasts, osteoblasts, myoblasts, neuroblasts,fibroblasts, glioblasts; germ cells, hepatocytes, chondrocytes,keratinocytes, cardiac muscle cells, connective tissue cells, epithelialcells, endothelial cells, hormone-secreting cells, cells of the immunesystem, neurons, and the like.

Examples also include cells that have been genetically engineered,transformed cells, and immortalized cells. One example of geneticallyengineered cells useful in the present invention is a geneticallyengineered cell that makes and secretes one or more desired molecules.When matrices comprising genetically engineered cells are implanted inan organism, the molecules produced can produce a local effect or asystemic effect, and can include the molecules identified above aspossible substances.

Cells may produce substances that inhibit or stimulate inflammation;facilitate healing; resist immunorejection; provide hormone replacement;replace neurotransmitters; inhibit or destroy cancer cells; promote cellgrowth; inhibit or stimulate formation of blood vessels; augment tissue;and to supplement or replace the following tissue, neurons, skin,synovial fluid, tendons, cartilage, ligaments, bone, muscle, organs,dura, blood vessels, bone marrow, and extracellular matrix.

The shape of the extracellular matrix may help send signals to the cellsto grow and reproduce in a specific type of desired way. Other factorsand differentiation inducers may be added to the matrix to promotespecific types of cell growth.

Once the tissue has been reduced to a suspension of individual cells,the suspension can be fractionated into subpopulations from which thecells elements can be obtained. This also may be accomplished usingstandard techniques for cell separation including, but not limited to,cloning and selection of specific cell types, selective destruction ofunwanted cells (negative selection), separation based upon differentialcell agglutinability in the mixed population, freeze-thaw procedures,differential adherence properties of the cells in the mixed population,filtration, conventional and zonal centrifugation, centrifugalelutriation (counterstreaming centrifugation), unit gravity separation,countercurrent distribution, electrophoresis and fluorescence-activatedcell sorting (see e.g. Freshney, (1987) Culture of Animal Cells. AManual of Basic Techniques, 2d Ed., A. R. Liss, Inc., New York, Ch. 11and 12, pp. 137-168). For example, salivary cells may be enriched byfluorescence-activated cell sorting. Magnetic sorting may also be used.The cells may also be isolated using magnetic selection techniques.

Cell fractionation may also be desirable, for example, when the donorhas diseases such as cancer or tumor. A cell population may be sorted toseparate the cancer or tumor cells from normal noncancerous cells. Thenormal noncancerous cells, isolated from one or more sorting techniques,may then be used for tissue reconstruction.

Isolated cells can be cultured in vitro to increase the number of cellsavailable for seeding into the biocompatible substrate. To prevent animmunological response after implantation of the artificial gift or limbcomposite tissue construct, the subject may be treated withimmunosuppressive agents such as, cyclosporin or FK506.

Isolated cells may be transfected with a nucleic acid sequence. Usefulnucleic acid sequences may be, for example, genetic sequences whichreduce or eliminate an immune response in the host. For example, theexpression of cell surface antigens such as class I and class Hhistocompatibility antigens may be suppressed. In addition, transfectioncould also be used for gene delivery. Cells may be transfected withspecific genes prior to seeding onto the biocompatible substitute. Thus,the cultured cells can be engineered to express gene products that wouldproduce a desired protein that helps ameliorate a particular disorder.

The tissue cells grown on the matrix substrate may be geneticallyengineered to produce gene products beneficial to implantation, e.g.,anti-inflammatory factors, e.g., anti-GM-CSF, anti-TNF, anti-IL-1, andanti-IL-2. Alternatively, the tissue cells may be genetically engineeredto “knock out” expression of native gene products that promoteinflammation, e.g., GM-CSF, TNF, IL-1, IL-2, or “knock out” expressionof MHC in order to lower the risk of rejection.

Methods for genetically engineering cells for example with retroviralvectors, adenoviral vectors, adeno-associated viral vectors,polyethylene glycol, or other methods known to those skilled in the artcan be used. These include using expression vectors which transport andexpress nucleic acid molecules in the cells. (See Geoddel; GeneExpression Technology: Methods in Enzymology 185, Academic Press, SanDiego, Calif. (1990).

Vector DNA is introduced into prokaryotic or eukaryotic cells viaconventional transformation or transfection techniques. Suitable methodsfor transforming or transfecting host cells can be found in Sambrook etal. Molecular Cloning: A Laboratory Manual, 2nd Edition, Cold SpringHarbor Laboratory press (1989), and other laboratory textbooks.

Once seeded onto the matrix, the cells will proliferate and develop onthe matrix to form a tissue layer. Importantly, because the matrix hasan infra-structure that permits culture medium to reach the tissuelayer, the cell population continues to grow, divide, and remainfunctionally active to develop into a tissue that has a morphology whichresembles the analogous structure in vivo.

It is important to recreate, in culture, the cellular microenvironmentfound in vivo for the particular tissue being engineered. By using amatrix that retains an infra-structure that is similar or the same as anin vivo tissue structure, the optimum environment for cell-cellinteractions, development and differentiation of cell populations, iscreated.

Growth factors and regulatory factors can be added to the media toenhance, alter or modulate proliferation and cell maturation anddifferentiation in the cultures. The growth and activity of cells inculture can be affected by a variety of growth factors such as growthhormone, somatomedins, colony stimulating factors, erythropoietin,epidermal growth factor, hepatic erythropoietic factor (hepatopoietin),and like. Other factors which regulate proliferation and/ordifferentiation include prostaglandins, interleukins, andnaturally-occurring chalones.

The artificial tissue constructs of the invention can be used in avariety of applications. For example, the artificial tissue constructscan be implanted into a subject to replace or augment existing tissue.The subject can be monitored after implantation of the artificial tissueor organ, for amelioration of the disorder.

Other embodiments and used of the invention will be apparent to thoseskilled in the art from consideration of the specification and practiceof the invention disclosed herein. All U.S. Patents and other referencesnoted herein for whatever reason are specifically incorporated byreference. The specification and examples should be considered exemplaryonly with the true scope and spirit of the invention indicated by thefollowing claims.

EXAMPLES Example 1 Methods and Materials

(i) Scaffold Preparation

Electrospun nanofiber scaffolds have been developed using a solution ofcollagen type I, elastin, and poly(D,L-lactide-co-glycolide) (PLGA, mol.ratio 50:50, Mw 110,000) (Boeringer-Ingelheim, Germany). Collagen type Ifrom calf skin (Elastin Products Company, Owensville, Mo.), elastin fromligamentum nuchae (bovine neck ligament), (Elastin Products Company,Owensville, Mo.), and PLGA are mixed at a relative concentration byweight of 45% collagen, 40% PLGA, and 15% elastin. The solutes aredissolved in 1,1,1,3,3,3-hexafluoro-2-propanol (99+ %) (Sigma ChemicalCompany, St. Louis, Mo.) at a total solution concentration of 15 w/v %(150 mg/mL). High molecular weight PLGA, previously used forelectrospinning tissue scaffolds is added to the solution to increasemechanical strength of the scaffold and increase viscosity and spinningcharacteristics of the solution.

Physically, the electrospinning method requires a high voltage powersupply, a syringe pump, a polymer solution or melt to be spun, and agrounded collection surface. During electrospinning, the groundedmandrel rotates while the stage translates to ensure even deposition offibers onto the mandrel surface. Solutions were electrospun using a highvoltage power supply (Spellman High Voltage, Hauppauge, N.Y.) at 25 kVpotential between the solution tip and the grounded surface. Thesolution was delivered with a 5 mL syringe through an 18 gauge blunt tipneedle at a flow rate of 3.0 mL/hr using a syringe pump. Fibers collectonto a grounded mandrel at a distance of 15 cm from the tip. The mandrelis a 303 stainless steel rod which is rotated at ˜500 rpm. The mandrelsize is initially 4.75 mm to allow for contraction of the graft due tocrosslinking. Uniform scaffolds of 120 mm length were created using 2.4mL of solution. This apparatus is shown schematically in FIG. 1.

Scaffolds were further crosslinked for increased stability and strength,using two crosslinking methods. The scaffolds were soaked for twominutes in 20% dextran solution in phosphate buffered saline prior tocrosslinking to reduce hydration-induced swelling and contraction of thescaffold. The scaffolds were crosslinked by immersion in 1) 1%glutaraldehyde solution and 2) EDC/NHS in MES/EtOH solution for 2 hoursat room temperature. These data show that it is possible to fabricatevascular scaffolds from biological polymers with mechanics and structuresimilar to decellularized scaffolds and native arteries.

FIG. 1. shows the electrospinning apparatus in which fibers deposit ontoa grounded collection surface as solvent evaporates due to increasingsurface area/volume ratio of solution. The electrostatic field causessplaying of solution, and solutions of sufficient viscosity and surfacetension form fibrous mats which adhere to grounded surfaces.

(ii) Cell Seeding

A confluent monolayer of endothelial cells is the most important barrieragainst thrombus formation, and endothelial cell mediated NO productionis important to maintain vascular tone. Cells were seeded with a mouseendothelial cell line MS 1 cells. The cells routinely cultured in tissueculture polystyrene flasks at 37° C. under 5% CO2 atmosphere wereharvested after the treatment with 0.1% trypsin-EDTA. The scaffolds weremounted in tissue culture dishes. After equilibration with PBS, thecells (1×10⁵ /mL) were seeded to the scaffolds. The culture medium usedwas DMEM medium containing 10% FBS, and antibiotics. After 2 daysculture, the cell attachment was assessed using scanning electronmicroscopy.

(iii) Microscopy

The relative quantity and distribution of collagen and elastin in avascular scaffold is important to the mechanical properties and functionof the seeded graft. To determine the distribution of components of thescaffolds, histo- and immunohistochemical analyses were performed toidentify collagen and elastin distribution.

(iv) Biocompatibility Testing (Cell Viability and Proliferation)

Long-term viability of cells is necessary for the seeded scaffold toremodel itself into a viable, patent vessel. Standard methods wereemployed to assess viability and proliferation. To test for cellviability, constructs were placed in 24-well plates with approximately100 mg of material per well. Four different types of material weretested for biocompatibility and cell survival, with one negative controlwell with no material: (1) GA-NFS (1% glutaraldehyde crosslinkedelectrospun scaffold); (2) EDC-NFS (EDC-crosslinked electrospunscaffold); (3) nBV (natural blood vessel, decellularized); (4) Latex(latex rubber, positive control).

Endothelial cells were seeded in the wells on a scaffold for testing viathe direct contact method. For cell viability, cell layers were rinsedwith PBS. 0.005% w/v neutral red was added in culture medium. Theneutral red solution was removed after 4 hours incubation at 37° C. with1% acetic acid and 50% ethanol solution by volume was added for dyeextraction, and dye extraction was shaken for 5 minutes. Absorbance wasthen measured at 540 nm using a spectrophotometer. The intensity of redcolor obtained was directly proportional to the viability of the cellsand inversely proportional to the toxicity of the material.

Cell proliferation was tested using the mitochondrial metabolic activityassay. Cell layers were first rinsed with PBS. MTT solution was added at1 mg/mL in PBS containing 1 mg/mL glucose. MTT solution was removedafter 4 hours incubation at 37° C. Dimethyl sulfoxide (DMSO) was used todissolve insoluble formazan crystals, and the absorbance at 540 nm wasmeasured using a spectrophotometer. The intensity of blue color wasdirectly proportional to the metabolic activity of the cell populationsand inversely proportional to the toxicity of the material or extract.

(v) Mechanical Testing

Compliance mismatch is one of the most common causes of vascular graftfailure, resulting in intimal hyperplasia and occlusion. If the scaffoldis too compliant, it may form an aneurysm.

Scaffolds were immersed in a water bath and cannulated at either end.One cannula was connected to a column of water and the other to adrainage tube. The column of water was high enough to create a pressurewithin the vessel-shaped scaffold of 120 mmHg. Water was drained throughthe scaffold in order to lower the pressure in increments of 10 mmHg. Ateach increment, the diameter of the scaffold was recorded using adigital camera. This process was repeated until the pressure was 0 mmHg.

(vi) Axial and Circumferential Segment Testing

Vessels must resist higher stress in the circumferential direction thanin the axial direction. Native vessels adapt their mechanics to thisloading environment. It is important that the electrospun scaffoldsexhibit a mechanical strength at least that of native vessels. Weperformed mechanical loading tests on the electrospun vessels in theaxial and circumferential directions using a uniaxial load test machine(Instron Corporation, Issaquah, Wash.). A short segment from a tubularscaffold was clamped at its cut ends for the axial test. The crossheadspeed was set at 0.5 mm/sec and the test was stopped when the straindecreased by 10% after the onset of failure. For testing in thecircumferential direction, a ring of material was cut from the scaffold,opened into a strip and then clamped at either end of the strip. Thistest was also performed at a rate of 0.5 mm/sec.

(vii) Burst Pressure Testing

The burst pressure for vascular scaffolds was found by monitoringincreasing pressures within the vessel until failure occurred. Apressure catheter was inserted through a cannulating fixture at one endof the vessel. A 60 cc pressure syringe was inserted through a customcannula at the other end of the vessel. The pressure was increased untilfailure or leakage occurred and the pressure change was recorded.

(viii) Functionalization of Matrices

To functionalize a matrix, EDC (10 mg) and sulfo-NHS (2 mg) were addedto 5 mL (0.05 mg/mL) of carboxylated quantum dots in aqueous solutionunder gentle stirring for 1 hr at room temperature. EDC activatedheparin (30 mg/20 μl) was prepared according to the same EDC and NHSmethod. In order to conjugate quantum dots and heparin, 5 mg PDA wasadded to the activated quantum dots and heparin solutions under stirringfor 2 hr at room temperature. The quantum dot-heparin (QD-heparin)conjugation can be quenched by adding an equal volume of 1 M Tris buffersolution (pH 7.4) and stored in 4° C. (FIG. 2).

(ix) Encapsulation

Microencapsulation of QD-heparin was performed by double emersion.Briefly, 4 mL of internal aqueous phase containing 30 mg QD-heparin and10 mg bovine serum albumin (BSA) as stabilizer was emulsified in 8 mlsolution of 100 mg PLGA and 100 mg PCL in dichloromethane. The solutionwas emulsified by vortexing for 5 minutes at room temperature. This W/Odispersion was diluted to 200 ml of 1% (w/v) aqueous PVA solution understirring for 4 hr at room temperature. The microcapsules were washedseveral times with deionized water and then lyophilized overnight.

(x) Heparin Release using IR Irradiation of the Quantum Dot

In order to evaluate the burst release of heparin, 0.55 mg of PLGAmicrocapsules containing QD-heparin were suspended in 2 ml of PBS(phosphate buffered saline). The solution was irradiated for 0, 10 and30 min using an AM 1.5 solar simulator at 75 mW/cm2. On days 1, 3 and 5,the samples were then cooled to 4° C., centrifuged at 4500 rpm for 20min and filtered (0.45 μm pore size) to remove any microcapsules for theoptical measurements. Luminescence measurements were performed using anargon ion laser (514.5 nm at 400 mW/cm2) as the excitation source andspectra were collected using a CCD spectrophotometer with an integrationtime of 40 sec.

(xi) Mouse Model

Mice (C57BL6) will be obtained from Jackson laboratories, Bar Harbor Me.All experimentation in mice will be performed aseptically under generalanesthesia (ketamine; 45-75 mg/kg and Xylazine; 10-20 mg/kg, IP). Theincision sites are scrubbed with betadine and wiped with alcohol.Analgesia (Buprenorphine 0.05-0.1 mg/kg, SC) is given post-operativelyafter implantation. Prophylactic antibiotic agents (cefazoline 25 mg/kg,sc) are given to the animals at the time of implantation. The preparedblood vessels (2×0.5 cm) will be implanted in the dorsal subcutaneousspace of mice through a minimal longitudinal midline incision with 2implants per animal. The wound will be closed with interruptedabsorbable sutures and the animals will be sacrificed 1, 2, 4, 8, 12, 18and 24 weeks after implantation for analyses. For the collection ofblood samples, mice will be anesthetized and blood will be retrievedinto heparin containing tubes using cardiac puncture and the mice willbe sacrificed thereafter.

(xii) Sheep Model

A total of 120 sheep will be used. The experimental study will consistof 6 different groups of the blood vessels. Each animal will serve asits own control. Animals will be sacrificed at 1, 3, 6, 12, and 18months after implantation. Animals will be monitored at 0, 1, 2, 3, and4 weeks and monthly for grafts implanted greater than one month.

Sheep will be sedated with Ketamine (5 mglkg, IM), intubated andanesthetized with Isofluorane (1-3%), and placed on a ventilatoradministering Isoflurane for maintenance. Following Duplex ultrasoundimaging of native femoral arteries the groins will be prepped in asterile fashion and antibiotics administered (cefazolin 25 mg/kg, i.v.).A longitudinal incision will be made overlying the superficial femoralartery, which will then be exposed over a length of 6 to 8 cm. Animalswill receive aspirin for 48 hours prior to surgery (80 mg, p.o.) andheparin will be administered immediately prior to implantation (100U/Kg, i.v.). The femoral artery will then be clamped and dividedproximally and an end-to-side anastomosis created between native andengineered artery with running 7-0 Prolene sutures. The distalanastomosis will then be created in a similar fashion and blood flowrestored through the implant. Duplex ultrasound will then be repeatedusing a sterile intraoperative probe cover to establish arterydimensions and blood flow immediately after implantation. Wounds willthen be closed with absorbable sutures and the animals recovered fromanesthesia using Atropine (0.02 mg/kg i.v.) prior return to standardhousing. Post-operative antibiotics will be administered (Cephazoline 25mg/kg/day) for 3 days following the procedure. Analgesia will beadministered (ketoprofen 2 mg/kg) every 6-12 hours for 3 days. Aspirinwill also be administered (80 mg daily) for 7 days orally foranticoagulation. The animals will be sacrificed 1, 3, 6, 12 and 18months after implantation for analyses. At each time point, 6 animalswill be euthanized for analysis.

Example 2 Preparation of Decellularized Organs

The following method describes a process for removing the entirecellular content of an organ or tissue without destroying the complexthree-dimensional infra-structure of the organ or tissue. An organ, e.g.a liver, was surgically removed from a C7 black mouse using standardtechniques for tissue removal. The liver was placed in a flaskcontaining a suitable volume of distilled water to cover the isolatedliver. A magnetic stir plate and magnetic stirrer were used to rotatethe isolated liver in the distilled water at a suitable speed for 24-48hours at 4° C. This process removes the cellular debris and cellmembrane surrounding the isolated liver.

After this first removal step, the distilled water was replaced with a0.05% ammonium hydroxide solution containing 0.5% Triton X-100. Theliver was rotated in this solution for 72 hours at 4° C. using amagnetic stir plate and magnetic stirrer. This alkaline solutionsolubilized the nuclear and cytoplasmic components of the isolatedliver. The detergent Triton X-100, was used to remove the nuclearcomponents of the liver, while the ammonium hydroxide solution was usedto lyse the cell membrane and cytoplasmic proteins of the isolatedliver.

The isolated liver was then washed with distilled water for 24-48 hoursat 4° C. using a magnetic stir plate and magnetic stirrer. After thiswashing step, removal of cellular components from the isolated wasconfirmed by histological analysis of a small piece of the liver. Ifnecessary, the isolated kidney was again treated with the ammoniumhydroxide solution containing Triton X-100 until the entire cellularcontent of the isolated liver was removed. After removal of thesolubilized components, a collagenous three-dimensional framework in theshape of the isolated liver was produced.

This decellularized liver was equilibrated with 1×phosphate buffersolution (PBS) by rotating the decellularized liver overnight at 4° C.using a magnetic stir plate and magnetic stirrer. After equilibration,the decellularized liver was lyophilized overnight under vacuum. Thelyophilized liver was sterilized for 72 hours using ethylene oxide gas.After sterilization, the decellularized liver was either usedimmediately, or stored at 4° C. or at room temperature until required.Stored organs were equilibrated in the tissue culture medium overnightat 4° C. prior to seeding with cultured cells.

Example 3 Electrospun Matrices

An electrospun matrix was formed using the methods outlined inExample 1. A solution of collagen type I, elastin, and PLGA, were used.The collagen type I, elastin, and PLGA were mixed at a relativeconcentration by weight of 45% collagen, 40% PLGA, and 15% elastin.

The resulting fibrous scaffold had a length of 12 cm with a thickness of1 mm. A 2 cm representative sample is depicted in FIG. 3. Thisdemonstrates the feasibility of spinning Type I Collagen and elastininto fibers from nanometer to micrometer diameter using concentrationsfrom 3% to 8% by weight in solution. These results also show that byadding PLGA (Mw 110,000) to the mixture, solutions with higher viscosityand improved spinning characteristics could attained. By increasing thesolution concentration to 15%, thicker, stronger scaffolds were able tobe built while maintaining the collagen and elastin components.

Collagen type I stained positively on the decellularized scaffolds,demonstrating uniform distribution. Elastin distribution within thescaffolds was determined by Movat staining. The electrospun scaffoldswith 15% elastin demonstrated a uniform elastin matrix throughout thescaffold wall. These findings indicate that the matrix content anddistribution of the electrospun scaffolds can be manipulated to achievevarious matrix compositions depending on the need.

Results of biocompatibility assays were calculated as a percentage ofnegative control and both electrospun scaffolds performed similarly tothe decellularized blood vessel. These data suggest that thebiocompatibility of electrospun scaffolds is similar to that ofdecellularized scaffold.

Results of mechanical testing for compliance show a typicalpressure-diameter curve for native vessels, as well as fordecellularized and electrospun scaffolds. The diameter change wasapproximately 5% for native vessels and electrospun scaffolds within thephysiologic pressure range which is consistent with the in vivomechanical behavior of porcine and human arteries (FIG. 4). These datademonstrate that the electrospun scaffolds created have a compliancesimilar to that of a native vessel.

Results of the axial and circumferential mechanical tests fromelectrospun scaffolds tended to exhibit a more isotropic behavior.Strain in the axial and circumferential directions were nearlyequivalent before failure occurred.

The results of burst pressure testing show that the burst pressure forthe electrospun construct was 1,425 mmHg or nearly 12 times systolicpressure. These data suggest that electrospun scaffolds have adequateinitial strength and elasticity to withstand the mechanical environmentwhen being surgically placed in the circulatory environment.

Histological analysis of the explanted vascular scaffolds from miceshowed that there was no evidence of inflammation or tissueencapsulation.

Collectively, these results show that it is possible to control thecomposition of electrospun scaffolds for use as vascular grafts. Higherconcentrations of collagen type I and elastin than previously employed,and mixing with PLGA, result in improved spinning characteristics andstrength of grafts, which resist almost 12× systolic pressure. Scaffoldsalso exhibited compliance characteristics similar to native arteries.Scaffolds had an average fiber diameter of 720 nanometers. EDCcrosslinked scaffolds demonstrate superior cell proliferationcharacteristics to glutaraldehyde crosslinked scaffolds as assessed bymitochondrial metabolic activity assay. Cell viability assays did notdemonstrate as pronounced a difference in crosslinking method. Theseresults are some of the first data on biocompatibility of electrospunscaffolds created with biological polymers and PLGA. This workdemonstrates the promise of electrospinning as a fabrication process forvascular graft scaffolds.

Example 4 Cross-Linking of Electrospun Matrices

This example demonstrates how to increase the strength and stability ofthe electrospun scaffold by chemical cross-linking. The scaffolds weresoaked in 20% dextran solution in phosphate buffered saline prior tocrosslinking to reduce hydration-induced swelling and contraction of thescaffold. The scaffolds were crosslinked by immersion in EDC/NHS inMES/EtOH solution for 2 hours at room temperature. Scanning electronmicrographs of the resulting fibers showed fiber diameters of 500 nm orless and a random orientation of fibers. Atomic force microscopy of thescaffold and a confocal image of nanofibers with an adhering endothelialcell demonstrate the scaffold structure. These data show that it ispossible to fabricate vascular scaffolds from biological polymers withmechanics and structure similar to decellularized scaffolds and nativearteries.

Example 5 Distribution of Collagen and Elastin Content

The relative quantity and distribution of collagen and elastin in avascular scaffold is important to the mechanical properties and functionof the seeded graft. The scaffold composition was assessed usinghistochemical analysis for collagen types I, II, and III, elastin andhematoxylin, and eosin (H&E) staining was also performed.

The levels of collagen type I, II, and III, and elastin fordecellularized matrices and collagen type I and elastin for electrospunmatrices were analyzed using computerized histomorphometric analysis.NIH Image/J Image analysis software (National Institutes of Health,Bethesda, Md.) was used for the analysis.

Immunohistochemical analyses using antibodies specific to collagen typesI, II and III were performed on the decellularized and electrospunscaffolds. The decellularized scaffolds showed similar collagen type Iand III in the vascular media, which corresponds to normal bloodvessels. In this study, 45% collagen type I was used to demonstrate thecontrollability of the scaffold fabrication. Collagen type I stainedpositively on the decellularized scaffolds, however, collagen type IIIstained negatively. Elastin distribution within the scaffolds wasdetermined by Movat staining. Abundant elastin fibers were observed inthe entire decellularized scaffold wall with a prominent distribution inthe serosal and luminal surface. The electrospun scaffolds with 15%elastin demonstrated a uniform elastin matrix throughout the scaffoldwall. These findings indicate that decellularized vascular scaffoldspossess matrices similar to normal vessels and that the matrix contentand distribution of the electrospun scaffolds can be manipulated toachieve various matrix compositions depending on the need.

Histograms of the distribution of color were used to determine relativeamounts of each component from each stain against negative controls. Allvalues were normalized by area for comparison. Amounts of collagen I,elastin, and PLGA were known for electrospun matrices because offabrication parameters. Calibrating the image data for relative amountsof collagen utilized both the normalized areas with negative controls,and was calibrated based on known composition of electrospun matrices.

The results demonstrate the composition of collagen I, II, and III, andelastin, in the decellularized scaffolds as well as componentpercentages in electrospun matrices. These studies show that thecollagen and elastin content of decellularized and electrospun scaffoldsis similar to that of native vessels.

Example 6 Compliance Testing of Scaffolds

Compliance mismatch is one of the most common causes of vascular graftfailure, resulting in intimal hyperplasia and occlusion. If the scaffoldis too compliant, it may form an aneurysm. This example describes how totest for compliance of the scaffolds. Decellularized and electrospunvessel shaped scaffolds were immersed in a water bath and cannulated ateither end. One cannula was connected to a column of water and the otherto a drainage tube. The column of water was high enough to create apressure within the vesselshaped scaffold of 120 mmHg. Water was drainedthrough the scaffold in order to lower the pressure in increments of 10mmHg. At each increment, the diameter of the scaffold was recorded usinga digital camera. This process was repeated until the pressure was 0mmHg. Results show the typical pressure-diameter curve for nativevessels, and the experimental curves for decellularized and electrospunscaffolds. The diameter change was approximately 5% for native andelectrospun and 15% for decellularized scaffolds within the physiologicpressure range which is consistent with the in vivo mechanical behaviorof porcine and human arteries. Thus, both decellularized and electrospunscaffolds have a compliance similar to that of a native vessel.

Example 7 Circumferential and Axial Loading of Decellularized andElectrospun Vessels

Vessels must resist higher stress in the circumferential direction thanin the axial direction. Native vessels adapt their mechanics to thisloading environment. It is important that the decellularized andelectrospun scaffolds exhibit a mechanical behavior similar to nativevessels. Thus, mechanical loading tests were performed on thedecellularized vessels and electrospun vessels in the axial andcircumferential directions using a uniaxial load test machine (InstronCorporation, Issaquah, Wash.). An entire vessel-shaped scaffold wasclamped at its cut ends for the axial test. The crosshead speed was setat 0.5 mm/sec and the test was stopped when the strain decreased by 10%after the onset of failure. For testing in the circumferentialdirection, a ring of material was cut from the scaffold, opened into astrip and then clamped at either end of the strip. This test was alsoperformed at a rate of 0.5 mm/sec. Results of the axial andcircumferential mechanical tests from electrospun scaffolds are shown inFIGS. 5A and 5B, respectively.

The electrospun scaffolds tended to exhibit a more isotropic behavior.Strain in the axial and circumferential directions were nearlyequivalent before failure occurred. In general, the decellularizedconstruct exhibits the orthotropic mechanical behavior that is expectedfrom the known mechanical behavior of arteries. In particular, strain inthe circumferential direction is lower than strain in the axialdirection. This was true for scaffolds prior to and after implantation.

The burst pressure for vascular scaffolds was found by monitoringincreasing pressures within the vessel until failure occurred. Apressure catheter was inserted through a cannulating fixture at one endof the vessel. A 60 cc pressure syringe was inserted through a customcannula at the other end of the vessel. The pressure was increased untilfailure or leakage occurred and the pressure change was recorded. Theresults show that the burst pressure for the decellularized constructwas 1,960 mm Hg or approximately 16 times systolic pressure. The burstpressure for the electrospun construct was 1,425 mm Hg or nearly 12times systolic pressure. We demonstrated that both electrospun anddecellularized scaffolds had adequate strength and elasticity and may besubstitutes for native vessels.

Example 8 Isolation, Characterization and Vessel Seeding of SheepProgenitor EPC and MPC

Progenitor EPC and progenitor muscle cells (MPC) were isolated from 60ml peripheral blood of the internal jugular vein of sheep. TheLleukocyte fraction was obtained by centrifuging on a Histopaque densitygradient. Some of the cells were resuspended in medium and plated onfibronectin coated plates. At 24 hr intervals the floating cells weretransferred to new fibronectin coated plates. EPC were induced by growthin EGM-2 medium that contained VEGF and bFGF. The rest of the cells werecultured in the presence of 10 μM 5-Azacytidin for 24 hours. Thereafterfloating cells were transferred to a new fibronectin coated plate andcultured in myogenic medium (DMEM low glucose containing 20% fetalbovine serum, 10% Horse Serum, 1% Chick Embryo extract and 1%antibiotics) in order to induce MPC. EPC and MPC were cultured for 4-6weeks in order to assume differentiated morphology. Immunohistochemicalanalysis of EPC showed that most of the cells expressed VE cadherin andCD31 but not Desmin. However, MPC showed expression of Vimentin andDesmin but not of VE cadherin. The expression of these markers wasmaintained during culture in vitro. These results indicate that culturedEPC and MPC possess EC and muscle cell phenotype, respectively.

EPC were labeled by PKH 26 green fluorescent dye and MPC were labeled byPKH 27 red fluorescent dye. Labeled EPC and MPC were seeded on theluminal and the outer surfaces of decellularized vessel segments,respectively, in order to demonstrate the biocompatibility of thedecellularized vessel. After 7 days the presence of red and greenlabeled cells on the decellularized vessel was noted. In addition,seeded vessels were seeded with a suspension of red-labeled MPC andgreen labeled-EPC (5×10⁶ cells/ml) and cells were allowed to grow for 7days. The vessels were embedded in OCT media in order to obtain frozensections. The sections were stained with DAPI. To detect cell nuclei,sections were visualized using a fluorescent microscope. Data shows thatEPC were maintained on the luminal side of the scaffold and MPC on theserosal surface.

Example 9 Cell Attachment

A confluent monolayer of endothelial cells is the most important barrieragainst thrombosis formation. Endothelial cell mediated NO production isimportant in maintaining the vascular tone. To examine cell attachment,the decellularized and electrospun vessels were seeded with endothelialcells. Cell attachment was assessed using scanning electron microscopyof scaffolds seeded with a mouse endothelial cell line (MS1). SEMmicrographs reveal a confluent monolayer on the inner surface of boththe decellularized and electrospun vessels at 48 hours. These resultsindicate that endothelial cells form confluent monolayers ondecellularized and electrospun scaffolds.

Example 10 Biocompatibility (Cell Viability and Proliferation)

Long-term viability of cells is necessary for the seeded scaffold toremodel itself into a viable, patent vessel. To test for cell viability,decellularized and electrospun constructs were placed in 24-well plateswith approximately 100 mg of material per well. Four different types ofmaterial were tested for biocompatibility and cell survival, with onenegative control well with no material: (1) GA-NFS (1% glutaraldehydecrosslinked electrospun scaffold); (2) EDC-NFS (EDC-crosslinkedelectrospun scaffold); (3) nBV (natural blood vessel, decellularized);(4) Latex (latex rubber, positive control).

Endothelial cells were seeded in the wells on a scaffold for testing viathe direct contact method. For cell viability, cell layers were rinsedwith PBS. 0.005% w/v neutral red was added in culture medium. Theneutral red solution was removed after 4 hours incubation at 37° C. with1% acetic acid and 50% ethanol solution by volume was added for dyeextraction, and dye extraction was shaken for 5 minutes. Absorbance wasthen measured at 540 nm using a spectrophotometer. The intensity of redcolor obtained was directly proportional to the viability of the cellsand inversely proportional to the toxicity of the material. Results werereported as a percentage of negative control, and both electrospunscaffolds performed similarly to the decellularized blood vessel (FIG.6A).

Cell proliferation was tested using the mitochondrial metabolic activityassay. Cell layers were first rinsed with PBS. MTT solution was added at1 mg/mL in PBS containing 1 mg/mL glucose. MTT solution was removedafter 4 hours incubation at 37° C. Dimethyl sulfoxide (DMSO) was used todissolve insoluble formazan crystals, and the absorbance at 540spectrophotometer. The intensity of blue color was directly proportionalto the metabolic activity of the cell populations and inverselyproportional to the toxicity of the material or extract. Thegluataraldehyde treated matrices show more pronounced differences thanin proliferation assays, with EDC treated scaffolds being similar tonatural blood vessels (FIG. 6B).

Cell viability and proliferation testing was also performed to determinethe effects of various concentrations of gadolinium (Gd) on thescaffolds, on cell survival (FIG. 7). The tests revealed little effectof Gd levels on cell viability or survival. The results indicate thatboth scaffolds can promote cell growth and thus may be used for thebioengineering of vascular grafts.

Example 11 External Functionalization of Matrices

This example describes how to generate matrices with image enhancingagents and quantum dots. In particular, Gd-DPTA and quantum dotfunctionalization of an external scaffold. The scaffold can be anybiocompatible substrate, such as a synthetic PGA matrix, an electrospunmatrix, or a decellularized matrix. At present, no clinically availablevascular graft allows for noninvasive monitoring of the integration ofthe graft in vivo, nor does any graft incorporate anticoagulants intoits structure. A reliable method is needed to attach nanomaterials toscaffolds, e.g., vascular scaffolds, in order to increase functionality,in particular as a material marker and for anticoagulation. CarboxylatedGd and quantum dot (QD) materials were coupled to the surface of boththe decellularized and the electrospun scaffolds using an EDC/sulfo-NHSmethod. Any unreacted material was quenched and removed by rinsing thescaffold with 0.1 M Tris buffer. The liquid from the final washing wascolorless under UV elimination.

Under blacklight illumination the functionalized scaffold showsmulticolor fluorescence. Areas of red-orange emission are from thequantum dots. The pale white color, which is stronger in intensity thanthe control tissue, comes from the Gd containing material that canfluoresce with a pale blue color. The data shows that it is possible toincorporate heparin onto the surface of a scaffold. The scaffolds arealso able to bind Gd.

Example 12 Internal Functionalization of Matrices

This example describes the production of electrospun matrices with imageenhancing agents and therapeutic agents. In particular, Gd-DPTA and QDaddition to the internal electrospun scaffolds. Fabricating vascularscaffolds using electrospinning provides an opportunity to incorporateimage enhancing agents within the bulk material. Solutions were spunsuccessfully containing gadolinium diethylenetriamine pentacetic acid(Gd-DPTA) in HFP at a concentration of 15 mg/mL and with quantum dotsadded at a concentration of 8% by volume from a quantum dot solution of25.5 nmol/mL in toluene. No morphological change was noted in thescaffolds due to the addition of the Gd-DPTA or the QDs. These resultsshow that incorporating nanoparticles into the scaffolds has only aminimal effect on the morphology of the resulting structure.

Example 13 Matrices with Quantum Dots

This example describes how to couple therapeutic agents, such as heparinto the quantum dots (QD). Heparin is a potent anticoagulant agent. Toavoid systemic administration, a method is needed to control the releaseof heparin from the vascular scaffold and to bind the heparin to thescaffold. In this experiment, EDC (10 mg) and sulfo-NHS (2 mg) was addedinto the 5 mL (0.05 mg/mL) of carboxylated quantum dots in aqueoussolution under gentle stirring for 1 hr at room temperature. EDCactivated heparin (30 mg) was prepared according to the same EDC and NHSmethod as described above. In order to conjugate quantum dots andheparin, 5 mg phenylene diamine (PDA) was added to the activated quantumdots and heparin solutions while stirring for 2 hr at room temperature.The quantum dot-heparin (QD-heparin) conjugation can be quenched byadding an equal volume of 1 M Tris buffer solution (pH 7.4) and storedin 4° C.

Microencapsulation of QD-heparin was performed by double emersion.Briefly, 4 mL of internal aqueous phase containing 30 mg QD-heparinconjugation and 10 mg bovine serum albumin emulsified in 8 mL of asolution of 100 mg PLGA (MW; 110,000) and 100 mg PCL (MW; 110,000) inDCM. The solution was emulsified by vortexing for 5 min at roomtemperature. This W/O dispersion was diluted into 200 mL of 1% (w/v)aqueous PVA solution under stirring for 4 hr at room temperature. Themicrocapsules (MCs) were washed several times with deionized water andthen lyophilized overnight. QD-heparin nanocapsules (NC) wereincorporated into scaffolds by placing the functionalized vascularscaffold in 1 wt % PLL in PBS. Vascular scaffolds were immersed in thePLL-nanocapsule solution for 3-4 hours, and lyophilized beforesterilization with gamma irradiation.

A fluorescence image of an isolated microcapsule containing quantum dotsshows that the characteristic fluorescence from the quantum dots used inthis experiment is at 500 nm. The data show that it is possible to bindheparin to quantum dots and encapsulate the bound heparin in abiodegradable polymer, for attachment to the vascular scaffold.

Example 14 Release Kinetics of Heparin: In vitro Release of Heparin andBurst Release by Irradiation

To assess the effectiveness of quantum dots for controlled delivery ofheparin, the release kinetics of the drug was analyzed following anirradiation burst. In order to evaluate the burst release of heparin,0.55 mg of PLGA microcapsules with QD-heparin were suspended in 2 ml ofbuffered saline solution. The solutions were irradiated for 0.0, 10, and30 min using an AM1.5 solar simulator at 75 mW/cm². On days 1, 3, and 5the samples were then cooled to 4° C. and centrifuged at 4500 rpm for 20min. The solutions were filtered (0.45 m pore size) to remove anymicrocapsules for the optical measurements.

Luminescence measurements were performed using an argon ion laser (514.5nm at 400 mW/cm²) as the excitation source and spectra were collectedusing a CCD spectrophotometer with an integration time of 40 sec.Irradiated samples showed increased luminescence over time indicating a“burst effect”. The kinetic profile of heparin confirms that irradiationinduced the burst release out of functionalized microcapsules. Heparinrelease was monitored by optical analysis (FIG. 8A) and biochemicalanalysis (FIG. 8B). These results indicate that NIR can be used toinitiate the release of heparin from the QD-heparin microcapsules.

Normally, heparin is administered at the site of implantationimmediately following surgery to prevent acute thrombosis. Afterwards,heparin is administrated within the first week twice a day by injection.In order to improve the patient's compliance, heparin could beimmobilized in vascular scaffolds for extended period of time. However,the immobilization of heparin to the scaffolds results in a slow releaseof heparin which is not appropriate for thrombus prevention. Toaccelerate the burst release of heparin, near infrared (NIR) irradiationof the quantum dots bound to heparin can to be used to achieve thisgoal.

Example 15 Determination of the Remaining Heparin in Retrieved VesselImplants from Mice

To assess the effectiveness of heparin in an in vivo model, heparin mustbe evaluated after the explantation of the scaffold. The remainingheparin in functionalized blood vessels (heparin-QD) implanted in micewas determined by toluidine blue staining and most of the heparin isshown to have diffused out of the vessel two weeks after implantation.The heparin content was analyzed by a Rotachrome kit and the dataconfirms that very little heparin remains after two weeks. The data showthat the activity of heparin was successfully prolonged in the scaffoldbeyond its normal 1-2 hour half-life (FIG. 9).

The inflammatory response of quantum dots should be addressed forclinical applications. From the histological analysis of the explantedvascular scaffolds from mice, there was no evidence of inflammation ortissue encapsulation. The data indicate that conjugated heparin had onlya minimal inflammatory response.

Example 16 Evaluation of the Anti-Thrombogenic Properties of HeparinImmobilized Vessels

Although heparin is a powerful anticoagulant, it was important to verifythat this property still exists after immobilization. Two methods ofheparin binding were tested. Thirty milligrams of heparin was incubatedin 20 mM EDC and 10 mM sulfo-NHS in PBS for 2 hours at room temperature,and a 3 mm diameter decellularized scaffold was then immersed inheparin-EDC solution for 2 hours at room temperature. Aftercross-linking, the sample was rinsed in PBS several times to completelyremove residual EDC. Subsequently immobilization of heparin by physicaladsorption was performed using Poly(L-lysine) (PLL): The 3 mm diameterdecellularized scaffold was incubated in 2 mg/mL PLL solution for 2hours at room temperature. The PLL-adsorbed scaffold was immersed in 15mg/mL heparin solution for 1 hour at room temperature. Theanti-thrombogenic property of each method was evaluated using wholeblood from sheep by toluidine blue staining. Immediate coagulation wasobserved from the decellularized scaffold while no significant sign ofcoagulation was found from both EDC and PLL reacted decellularizedscaffolds 36 hours after blood treatment. The heparin-PLL decellularizedscaffold demonstrated the weakest staining which indicated the highestloading of heparin in the scaffold. These results showed thatimmobilized heparin was effective in preventing thrombus.

Example 17 Enhanced MRI Imaging

This example demonstrates the improved imaging observed with gadolinium.In vitro experiments were conducted on cell scaffolds with gadolinium todetermine the improvement in magnetic resonance imaging. Cylindricalcell scaffolds 20 millimeters long with an internal radius of 10millimeters and a outer radius of 14 millimeters were created withdifferent Gd loading concentrations. Cell scaffolds were individuallyplaced in test tubes and submerged in PBS. The four test tubes werearranged left to right in the following order: non-functionalized cellscaffold (control 1), functionalized scaffold (control 2), 1×Gdconcentration cell scaffold, 100×Gd concentration (control 3) and a1000×Gd concentration cell scaffold. (100×designates a concentration insolution during functionalization of 55 mg/kg of Gd-DPTA) Axial T1weighted spin echo images were acquired on a on a GE HealthcareTechnologies magnetic resonance imaging (MRI) 1.5T TwinSpeed scanner.

The T1 weighted image acquired with a phased array coil and a 200millisecond repetition time (TR) was obtained. Additional imagingparameters are as follows: echo time (TE)=13 ms, slice thickness=0.8 mm,256×128, field of view (FOV)=12 cm×6 cm, number of averages=100, andphase direction was right to left. The cell scaffolding loaded with1000×Gd (right most test tube) is clearly visible compared to controls1, 2, and 3. Samples were washed twice with TRIS buffer and PBS andstored in PBS for 2 weeks prior to imaging.

The previously described experiment was repeated for two different Gdloaded scaffold preparations: surface and volume loading. The scaffoldon the left is a cylindrically shaped scaffold identical to thepreviously described experiment with a surface loaded 1000×Gdpreparation. The scaffold on the right is a planar sheet of scaffoldwith the Gd embedded throughout the electrospun fibers as describedpreviously. The scaffold that had the Gd electrospun into the fibershowed a much higher contrast. The normalized signal intensities of thescaffold for the surface preparation and volume preparation are 1.5±0.2and 2.98±0.35, respectively. The data on MRI Imaging of Gd loadedscaffolds showed that Gd increases MRI contrast in proportion to thelevel of Gd loaded in the scaffold.

Example 18 In vivo Preliminary Data on Rodents

Although Gd may be maintained in the scaffolds in vitro, it is necessaryto demonstrate that it retains functionality in vivo. This experimentinvestigates the in vivo functionality of the scaffolds. Electrospunvascular scaffolds were implanted subcutaneously in a mouse for twoweeks prior to imaging. Gd was added to one of the vascular scaffolds toenhance its contrast on a T1 weighted image. A sagittal localizer imagewas acquired from the mouse and a T1 weighted coronal image containingthe two scaffolds was prescribed off the sagittal image. The importantimaging parameters of the T1 weighted image are repetition time (TR) 300milliseconds, echo time (TE) 14 milliseconds, and slice thickness 2millimeters. A 50% improvement in image contrast of the Gd scaffoldcompared to the control. These results in a rodent model demonstratethat the characteristics seen in vitro are maintained in vivo.

Example 19 Ex vivo Preliminary Data on Sheep Engineered Vessels

In vivo results in the rodent model were limited to subcutaneousspecimens. It was necessary to demonstrate similar results in a scaffoldexposed to blood flow in a large animal model. To determine thefeasibility of using the cell seeded scaffolds containing thenanoparticles (heparin conjugated with quantum dots and Gd-DTPA),femoral artery bypass procedures were performed in sheep. Peripheralblood samples were collected, circulating progenitor cells were selectedand differentiated into endothelial and smooth muscle cells in culture.Each cell type was grown, expanded separately and seeded ondecellularized vascular scaffolds containing the nanoparticles (30 mmlong). Nanoparticle containing scaffolds without cells served as acontrol. Under general anesthesia, sheep femoral arteries were imagedwith duplex ultrasonography (B-mode ultrasound and Doppler spectralanalysis) with a high resolution 15 MHz probe (HDI-5000, ATL) prior toscaffold implantation. The femoral artery was exposed through alongitudinal incision over a length of 6 to 8 cm. Aspirin and heparinwere used as anticoagulation and the femoral artery was clamped anddivided proximally. An end-to-side anastomosis was created betweennative and engineered artery. The distal anastomosis was created in asimilar fashion and blood flow restored through the implant followed byligation of native femoral artery between the two anastomoses. Dopplerultrasonography was performed using a sterile probe to establishscaffold dimensions and blood flow after implantation. Wounds wereclosed and the animals recovered from anesthesia prior to 3500 return tostandard housing. Aspirin was administered routinely for 7 days orallyfor anticoagulation.

Duplex ultrasound imaging was performed to determine the presence ofthrombosis, lumen narrowing intimal hyperplasia and graft wallstricture, and graft aneurismal degeneration. Longitudinal andcross-sectional images of the pre- and post operative arterial segmentsshowed a patent lumen 0 with similar peak systolic, end-diastolic andtime averaged velocities as the normal artery. The arterial wallthickness and luminal diameter of the engineered bypass was similar tonative artery. The engineered arterial bypass and the contralateralnormal femoral artery were scanned with MRI. T1 weighted spin echo MRimages were acquired with the following parameters: 256×126 matrix, 12×6mm FOV, 400 ms TR, 13 ms TE, 1 mm slice thickness, and 50 excitations.Average signal intensities of the samples were normalized by thebackground water intensity to account for receiver coil nonuniformities.The normalized intensities were 2.62 and 2.10 for the scaffold andnormal vessel, respectively.

This experiment was repeated at several different TRs and the signalintensity measured for the scaffold and the normal vessels. As expected,the signal intensity for the gadolinium enhanced scaffold is alwaysgreater than the normal vessel. These results confirmed that Gd andheparin loaded decellularized scaffolds maintain patency in a sheepmodel and maintain MRI contrast.

Gadolinium is a MR contrast agent that enhances images primarily bydecreasing the spin-lattice relaxation time (T1) of protons in tissues.Unlike radionuclides, it will remain effective as long as it islocalized in the engineered vessel. These results shown in vitro throughrepeated rinsing of the Gd doped scaffolds and in vivo through imagingof the engineered vessel, that the functionalized Gd nanoparticles arestable in the matrix. Within the first 3 months, approximately 80% ofthe graft will be remodeled. The Gd localized in the matrix willinitially enhance the imaging of the graft. The change in MR signal overtime, as the concentration of Gd decreases with remodeling of thevascular graft, will allow us to quantify the remodeling rates.

Example 20 Histomorphological Characteristics of Bypass Grafts in Sheep

To demonstrate cell attachment on the retrieved engineered vesselsinitially seeded with endothelial and muscle cells, scanning electronmicroscopy was performed 2 weeks after implantation. The implanteddecellularized scaffolds seeded with cells showed a uniform cellattachment on the luminal surface of the engineered artery similar tonormal vessels. The scaffolds without cells failed to exhibit cellattachment. These observations indicate that the cells seeded ondecellularized vascular scaffolds are able to survive and remainattached after surgery.

To assess the histo-morphological characteristics of the retrievedtissue from engineered arterial bypass grafts in sheep, histologicalevaluation was performed. The engineered arterial specimens were fixed,processed and stained with hematoxylin and eosin (H&E) and Movatstaining. The cell seeded engineered grafts contained uniformcellularity throughout the vascular walls. Abundant elastin fibers wereobserved in the entire arterial wall with a prominent distribution inthe serosa and luminal surface. These findings demonstrate that theengineered vessels, seeded with peripheral blood derived progenitorcells differentiated into endothelial and smooth muscle cells, are ableto show an adequate cellular architecture similar to native vessels.

Collectively, these studies show that it is possible to fabricate andfunctionalize both decellularized and electrospun scaffolds with cells(endothelial and smooth muscle) and nanomaterials (quantumdot—conjugated heparin) that are known to have a positive therapeuticbenefit. Moreover, the data shows the successful incorporation ofmolecules (gadolinium) enhancing MRI contrast to monitor the engineeredvessels over time. The combination of functionalization and imagingoffers the potential for making these scaffolds an ideal vascularsubstitute. The matrices are biocompatible, possess the ideal physicaland structural properties, and have been shown to be functional for over4 months in the carotid artery of sheep.

Example 21 To characterize the Engineered Vascular Grafts

In order for a vessel to function normally, it should have theappropriate structural properties to accommodate intermittent volumechanges. In pathologic conditions, normal vessel function and mechanicalproperties may be compromised. To translate the use of bioengineeredvessels to patients, it is first necessary to confirm that normalvessels are being formed, and that they retain adequate phenotypic andfunctional characteristics over time, especially with growth.

(i) Mechanical Testing

Understanding the mechanical properties of explanted vessels providesinformation about the adaptive remodeling those vessels have undergonewhile in the host animal. Mechanical testing will include arterialelongation (axial and circumferential), compliance, burst pressure,stress relaxation, and creep.

(ii) Phenotypic and Composition Analyses

Histological and immunohistochemical analysis can be performed on theretrieved vascular grafts. Longitudinal and cross sections will be takenfrom the transition zones between native vessels and graft and from therest of the graft. Specimens will be fixed, processed and stained withHematoxylin and eosin (H&E) and Masson's trichrome. Cross-sectionalareas of the adventitia, media, intima and lumen will be measured usingcomputer-assisted analysis of digital images (NIH Image Software). Inaddition to cross-sectional analysis of the engineered artery body, aseparate analysis will be performed for the anastomoses region betweennative and engineered arteries. The proximal and distal anastomoses willbe fixed in formalin, embedded in paraffin, and then cut incross-section for analysis of lumen caliber and artery wall thickeningin step-sections spanning each anastomosis. In parallel, quantitation ofthrombus formation will be performed using H&E staining. The phenotypiccharacteristics of the retrieved tissues will be determined over time.

To determine the degree of endothelial and smooth muscle content of thebioengineered vessels over time, in comparison to normal tissues,multiple molecular markers will be probed immunocytochemically and withWestern blot analyses, as described above. These markers will includeAnti-Desmin and Anti-Alpha Smooth Muscle Actin, which specificallydetects smooth muscle cells. Endothelialization will be evaluated byanti-von-Willebrand factor anti-CD-31 and anti-VEGF receptor, KDR,antibodies, which stain EC specifically. Cell proliferation andapoptosis in engineered arteries will be determined by BrdUincorporation and TUNEL staining.

The composition and distribution of extracellular matrix components,such as collagen and elastin, are important for the normal function ofblood vessels. While the collagen network is responsible for tensilestrength, elastin is important for the elastic recovery of the vessel.Therefore, an assessment of the collagen and elastin content anddistribution of the retrieved tissues over time will be performed withhistological and quantitative biochemical assays. To determine whetherthe retrieved vessels possess normal concentrations of collagen andelastin, as compared to normal controls, the total collagen and elastincontent per unit wet weight of the retrieved tissue samples will bemeasured quantitatively using the Sircol collagen and the Fastin elastinassay systems (Accurate Chemical & Scientific Corporation, Westbury,N.Y.). To determine the anatomical distribution of collagen within theengineered vessels, as compared to controls, Immunocytochemicallocalization of collagen types I, II and III will be performed usingspecific monoclonal antibodies (Southern Biotechnology Associates, Inc.,Birmingham, Ala.) and with the elastin-specific stain, Movat.

(iii). Physiological Analysis

The ability to synthesize vasoactive agents such as Nitric Oxide (NO)will further determine the functionality of the engineered vascularscaffolds. There is increasing evidence on the importance of NO invascular hemostasis. NO contributes to resting vascular tone, impairsplatelet activation, and prevents leukocyte adhesion to the endothelium.

Briefly, guinea pig thoracic aorta will be harvested, the endotheliumlayer removed by gentle rubbing and cut into 5-mm segments. Each segmentwill be suspended between 2 tungsten stirrups for measurement ofisometric tension. The vessel segments placed in an organ chamber with10 ml Kreb's buffer solution at 37° C. with a mixture of 5% CO₂, 15% O₂and a balance of N2. Each vessel (2-3 cm in length) is tied to a 21 Gneedle, which was attached to plastic IV tubing and placed above theorgan chamber with the fresh aortic segment. The segments will becontracted with 80 mM KCl Kreb's buffer in a stepwise fashion to obtaina resting tension of 4 g. After resting for 90 minutes, the segments arecontracted in response to prostaglandin F2α up to a final concentrationof 10-7M and until a stable contraction of approximately 50% of maximumKCl-induced contraction is achieved. Vasoactive agents and antagonistsare then added using an infusion pump through the vessels to induce NOproduction. Doses of the vasoactive agents between 10⁻⁷-10⁻³ M will betested and dose-response curves will be constructed.

Example 22 Creating an Artificial Digit

To demonstrate that a small limb can be restored in situ with multipletissue types, culture expanded bovine muscle progenitor cells (MPC) andosteoblasts or chondrocytes were seeded on their respective region ofcomposite scaffolds to engineer tissue in vivo. The cell seededscaffolds (“construct”), was implanted in the subcutaneous space ofathymic mice. The constructs maintained their volume and initialstructure when retrieved 2 months later. Grossly, the engineeredcomposite tissues formed rigid, cartilaginous or bony tissues that wereattached to the soft muscle tissue. Histochemistry,immunohistochemistry, and Western blots of the retrieved tissue withtheir specific markers confirmed muscle, cartilage, and bone formation.

To determine whether the engineered composite tissues possess adequatestructural and functional characteristics for their use in therestoration of limb tissue, the explanted composite tissues underwentmechanical and physiological testing. The compression-relaxationanalysis of the engineered cartilage tissues demonstrated preservationof structural integrity and resisted high compressive pressures.

Physiological organ bath studies of the retrieved muscle tissuesdemonstrate adequate contractility in response to electric fieldstimulation, indicating the presence of muscle function. These resultsindicate that a composite tissue consisting of rigid cartilage and/orbone and elastic functional muscle can be bioengineered and can be usedfor limb replacement.

For in vivo studies, the digit, composed of muscle, bone and tendon, canbe engineered in situ in a rabbit model for the restoration of a missinglimb.

Example 23 Creating an Intelligent Scaffold

Optimization of digit growth will include the development of new,“intelligent” biomaterial scaffolds that possess the necessaryultrastructural, biomechanical and biological characteristics requiredfor cell attachment, survival, neovascularization, innervation, andtissue maturation in an environment with limited blood supply (i.e. thefinger). To engineer a large functional limb tissue in situ, adequatevascularization and innervation are essential.

This example describes the creation of “intelligent” scaffolds thatrecruit progenitor cells and facilitate the formation of vascularsupport and innervation via controlled release of growth factors. Theintelligent scaffolds accelerate morphogenesis, tissue formation,maturation and function in extensively damaged limbs, to restore digitor limb function to the injured subject.

To engineer a functioning limb that would provide structural support andlocomotion (FIG. 10A), a composite scaffold system that accommodateseach cell type is essential (FIG. 10B). Scaffolds that deliver differentcell types have been created that allow a multiple range of motorfunction in order to demonstrate the feasibility of engineering afunctioning limb (FIG. 10C), which shows contraction, movement (bending,flexing) and relaxation. In the ex situ rodent study, polyglycolic acid(PGA) scaffolds were used to configure the rigid portion of the digit,seeded with either cartilage or bone cells, and a collagen based matrixderived from porcine bladders to allow movement by seeding with musclecells. The PGA scaffolds were composed of 2 cylindrical rods connectedto strips of collagen matrix fibers. This configuration allowed forcontraction, relaxation and movement of the composite engineereddigit-like structure

A similar approach was used to mimic native digit structure. Bonescaffolds will be fabricated with pulverized collagen mixed withpoly(D,L-lactide-co-glycolide) (PLGA) to achieve improved mechanical andbiocompatible properties. Such scaffolds have been fabricated andpossess the necessary structural morphology and surface chemistry forcell attachment and growth. Bone scaffolds can be tested forhydrophilicity, mechanical properties, cell viability, as well as celladhesion and growth.

For tendon and muscle scaffold fabrication, bladder-derived collagenfiber strips will be weaved and anchored to the bone scaffolds. Thesebiomaterials have adequate biomechanical, functional andbiocompatibility properties. The properties of each component of thedigit scaffolds will be further refined and optimized to achieveadequate structural and biomechanical properties for in situimplantation in the distal limb of the rabbit. Tests involving celladhesion, proliferation and differentiation will be performed. Physicalcharacteristics, including the hydrophilicity, porosity, degradability,and mechanical properties, of each component of the composite scaffoldswill be determined.

Example 24 To Examine the Use of Intelligent Scaffolds with Larger Limbs

The incorporation of vascular endothelial growth factor VEGF in theengineered muscle significantly enhanced vascularization and muscletissue mass in vivo. Another critical component involved in therestoration of functional limb tissue is neuronal innervation. Nervegrowth factor (NGF) is a potent axonal guiding mediator that promotesreinnervation of tissues. It has been demonstrated that axonalregeneration is significantly accelerated in the NGF contained scaffold,as compared to controls without NGF (FIG. 11). To assess thefunctionality of the intelligent scaffold, scaffolds will be seeded withrabbit myoblasts and bone cells, and this construct implantedsubcutaneously in athymic mice. The engineered muscle tissue will beretrieved at different time points for analysis. Neovascularization willbe evaluated by specific EC staining (CD31) and calculating vesselnumber and distribution. Innervation will be evaluated by specific nerveand acetylcholine receptor expression markers and counting the number ofaxonal processes and nerve endings using techniques described by Yiou,et al. (2003) Transplantation; 76: 1053-60. To develop an intelligentscaffold system that would release active growth factors for an extendedperiod in a controlled manner, technology involving drug delivery system(DDS) is necessary (Rafiti, et al. (1997) J Control Release; 43:89-102).VEGF and NGF will be incorporated into the composite digit scaffoldsusing in situ impregnation and microencapsulation methods. Growthfactors will be loaded into individual microspheres usingwater-in-oil-in-water (W/O/W) emulsion techniques.

In vitro release kinetics will be determined using enzyme-linkedimmunosorbent assay (ELISA). Biological activities of the growth factorswill be evaluated by cell (endothelial and nerve) proliferation assays.In vivo pharmacokinetic studies of the growth factors containedintelligent scaffolds will be performed in athymic mice. Degradation andlocal tissue response caused by the scaffolds will be tested.

Example 25 In situ Implantation of a Digit Composed of Bone, Muscle andTendon, in Rabbits

To determine the feasibility of replacing a missing limb, theregeneration of digit segments with an interconnecting joint of ananimal finger (phalanxes) can be performed. Interconnected-compositetissues can be engineered ex situ in the subcutaneous space of mice. Totest for in situ implantation, two interconnecting phalanxes of a rabbitdigit will be replaced using composite digit tissues composed of bone,muscle and tendon. Due to the limited tissue compartment and bloodsupply in this area compared with the subcutaneous space of mice, thegrowth and function of the composite tissue will be monitored.

The existing phalanxes will be excised and the engineered digit segmentsof the same length and caliber will be interposed and anastomosed. Theengineered digit will be evaluated in vivo prior to sacrifice atpre-determined time points. Achievement of digit function is the primegoal of limb restoration. To accurately assess the implanted engineeredtissue, in vivo structural and functional assessment will be necessary.Radiography, electromyography (sensory and motor function), physicalexamination (degree of mobility, extensor and flexor function) andmechanical testing (resistive, compressive and tensile force) will beperformed.

To demonstrate that the restored digit tissue can be used for a limbreplacement, the retrieved digits will be assessed by structural,mechanical and physiological testing. Mechanical testing will beperformed on the bone, tendon, muscle, tendon-bone interface andtendon-muscle interface of the explanted digit using a mechanical tester(Instron). The degree of resistive force in response to compression,compression-relaxation, tension and a multi-directional bending will betested to determine the levels of tissue maturation. The engineeredskeletal muscle tissue will be detached from the bone for musclefunctional evaluation. The engineered muscle will be activated usingboth receptor (i.e., nicotinic acetylcholine receptor) andnonreceptor-mediated (i.e., KCl depolarization) stimulation, as well aselectrical field stimulation, using standard physiologic organ bathtechniques. Statistical and computer analyses will be used to rigorouslyevaluate the physiological characteristics of the retrieved segment. Avariety of measures, including the rate, magnitude and duration ofcontractile and relaxation responses, as well as theelectrophysiological characteristics of the muscle will be used. In allcases, the retrieved skeletal muscle segments will be compared to the“native” tissue on the contralateral side. Tissue morphology,composition and organization will be assessed by scanning electronmicroscopy, histo- and immunohistochemistry and molecular analysis.

Example 26 Accelerated Tissue Organization and Maturation of SkeletalMuscle Using a Pre-Conditioning Bioreactor System

Muscle tissue organization and maturation is important in achievingmuscle tissue function in vivo. Bioreactor studies will be used todetermine the optimal conditions for enhancing unidirectional fiberorientation and accelerating muscle tissue maturation. This must beaccomplished prior to implanting the muscle tissue in the rabbit withthe other composite tissues (i.e., bone and tendon) so as to permit thein vivo environment to complete the maturation process toward fullyfunctional muscle. To achieve this goal, scaffolds seeded with musclecells will be placed in a bioreactor, and intermittent sinusoidalloading will be employed to provide the muscle with mechanical stimulus.

To this end, myoblasts will be expanded and seeded on muscle scaffoldsfor pre-conditioning in our bioreactor. The cell containing scaffoldswill be cycled at 10-25% of their initial length incrementally over thecourse of the incubation period. The cycling rate will also be increasedincrementally during the course of the bioreactor studies. Initialparameters for stretch and cycling rate will be selected with guidancefrom previous work in the field. (See e.g., Powell, et al. (2002) Am JPhysiol Cell Physiol;283:C1557-65; and Vandenburgh (2002) Ann NY AcadSci;961:201-2). However, determination of the optimal parameters forstretch and cycling rate will depend on the experimental results, andwill be selected based on the configuration that produces the mostcompatible physiological characteristics.

Example 27 To Demonstrate that an Entire Digit can be Engineered in situfor Functional Limb Restoration and Engineering of Multiple DigitSegments and in Situ Implantation in a Rabbit Model

A larger engineered limb tissue for replacement can be created. A tissuethat contains all the necessary tissue elements required to restore alimb, that is bone, muscle and tendon can be created as described above.To achieve a larger sized tissue, the improved intelligent scaffold andtissue maturation system will be used that would permit the enhancedfunctional tissue formation and maturation. A scaffold system that iscapable of delivering angiogenic and neurogenic factors in a controlledmanner will be fabricated and configured to serve as a multiple jointeddigit construct. Bone marrow-derived bone cells will be seeded on thebone scaffold and in vitro engineered organized muscle tissue will beattached to the scaffold.

The multiple jointed intelligent scaffolds containing the bone andmuscle cells will be tested in rabbits through an in situ implantation.The entire digit phalanxes will be excised from the rabbit and theengineered digit segments of the same length and caliber will beinterposed and anastomosed. The engineered digit segments will beevaluated in vivo prior to sacrifice at pre-determined time points.

The evaluation of the restored digit tissue will be performed in vivousing radiological analysis to determine bone formation,electromyography to determine muscle function and mechanical testing todetermine tissue strength. Radiographic evaluation will be performed todetermine the levels of bone, tendon and muscle formation, andelectromyography will be performed to demonstrate the degrees of motorand sensory innervation and contractile response of muscle tissue insitu. Physical examination will be performed to test joint and digitmobility, extensor and flexor function. Mechanical testing will also beperformed to determine the resistive, compressive and tensile forces ofthe engineered tissue in vivo.

Structural, biomechanical and physiological testing of the retrievedmultiple digit segments will also be performed. Structural morphologywill be assessed grossly (quantitatively) and by bone densitometry. Theengineered bone mineral content will be assessed, as well as initial andfinal dimensions of the muscular and tendinous components. In additionto the biomechanical testing methods described above, biomechanicaltesting of the joint will include bending and stiffness of the joint asmeasured in three-point bending, and joint dislocation strength measuredby failure in tension. Tissue morphology, composition and organizationwill be assessed by scanning electron microscopy, histo- andimmunohistochemistry and molecular analysis.

Example 28 Making and Using the Novel Composite Scaffolding SystemComprising a Biodegradable Synthetic Polymer and Naturally DerivedCollagen Matrix

This example demonstrates the successful fabrication of a bio-hybridbone scaffold composed of a biodegradable synthetic polymer and anaturally derived collagen matrix that possess necessary characteristicsfor bone tissue regeneration. In this example, BSM is used as thenaturally derived collagen matrix and PLGA is used as the biodegradablesynthetic polymer. The BSM-PLGA composite scaffolds are hydrophilic andpossess porous structures with a consistent interconnectivity throughoutthe entire scaffold which resulted in uniform cell seeding, adhesion andproliferation. The BSM-PLGA composite scaffolds are non-toxic, easilyfabricated and provide structural features that enhance the formation ofbone tissue for therapeutic regeneration.

Creation of bone tissue using cells requires a scaffold that serves as acell carrier which would provide structural support until bone tissueforms in vivo. The scaffold for bone tissue engineering should bebiocompatible and possess mechanical stability, a controlled degradationrate, hydrophilic surface chemistry and an appropriate porosity for cellaccommodation. In this invention, a composite scaffold for boneregeneration is disclosed that meets these criteria by hybridizing BSMas a natural bioactive material with synthetic PLGA polymers. Thisexample demonstrates that the BSM-PLGA composite scaffolds arenon-toxic, easily fabricated, and provide structural support withabundant pores with good interconnectivity.

BSM was selected for its biocompatibility, hydrophilic nature andability to induce cell proliferation. BSM consists mainly of type I andtype III collagen, elastin fibers and various proteins such asfibronectin and vitronectin that contain the arginine-glycine-aspartate(RGD) peptide binding motif. The presence of these extracellularproteins may enhance cell adhesion, survival and proliferation. However,BSM alone is well suited for bone graft applications due to its smallpore size, poor interconnectivity and inability to maintain structuralintegrity. To overcome these limitations, a natural-synthetic hybridscaffold was designed that would possess an interconnected network ofpores and sufficient mechanical and physicochemical properties thatwould maintain structural integrity, thus preventing collapse during thehandling and implantation process. The microarchitecture of thecomposite scaffolds did not differ from the control PLGA scaffold,indicating that the incorporation of BSM did not alter structuralchanges during the fabrication process. Further, the composite scaffoldsprovided desirable surface properties necessary for cell attachment andproliferation.

The surface properties of biomaterials play a major role in cellularinteraction, such as cell adhesion, infiltration and proliferation. Thesurface characteristics of many synthetic polymers including the PLGAare hydrophobic, which unfavorably influence their biocompatibilityduring the initial stage of contact with the biological environment. Asshown below, a hydrophilic surface was successfully achieved withoutcompromising the structural properties of the scaffold by introducingthe BSM component in the composite scaffold. The hydrophilicity of theBSM-PLGA composite scaffold was found to rely strongly on the content ofBSM, which mainly possesses hydrophilic proteins such as collagen andelastin. Therefore, enhanced penetration of culture medium, enhancedcell seeding and uniform distribution and accelerated cell migrationwere achieved.

A scaffold material for bone regeneration should be biocompatible andsafe for implantation. Cellular interactions of the BSM-PLGA compositescaffold were tested by using the two widely accepted complementaryassays; cell viability and mitochondrial metabolic activity. The cellviability assay using Neutral Red, which is based on dye incorporationinto lysosomes, and the MTT assay, which is based on the intact activityof a mitochondrial enzyme, demonstrated that the BSM-PLGA compositescaffolds are safe for the cells tested.

In this example, the composite scaffolds was tested with two differentcell types that could be used in orthopedic applications. Primary matureosteoblasts and stem cells form bone tissues when implanted in vivo ordifferentiated into bone cells. These cell types were used to determinetheir ability to survive, adhere and proliferate on the compositescaffolds. Cell accommodation, adhesion and proliferation wereapproximately 80% higher in the BSM-PLGA composite scaffolds, ascompared to the control scaffolds using these two cell types.Fabrication of the scaffolds with an appropriate pore size, porosity andsurface hydrophilicity resulted in abundant cell accommodation withincreased cell proliferation. The incorporation of BSM significantlyimproved the biological activities of the scaffold, while maintainingphysical and structural stability.

1. Materials and Methods

A. Materials

Porcine bladders, obtained from donor animals, were decellularized usinga multiple step process (De Filippo R E, Yoo J J, Atala A. J Urol 2002;168:1789-92). Briefly, the bladders were rinsed thoroughly withphosphate-buffered saline (PBS). The bladder submucosa wasmicrodissected and isolated from the muscular and serosal layers. Undercontinuous agitation, the bladder submucosa was rinsed with deionizedwater for 24 hr and placed in a solution containing 0.2% Triton X-100and 0.03% ammonium hydroxide for 14 days to remove all cellularcomponents. The bladder submucosa was washed with deionized watercontaining 10% cefazolin (APOTHECON, G. C. Hanford Mfg. Co., Syracuse,N.Y., USA) for an additional 24 hr. The BSM was freeze-dried at −50° C.for 48 hr using a lyophilizer (FreeZone 12, Model 775410, LABCONCO,Kansas City, Mo., USA), followed by pulverization in a freezer mill(SPEX 6700, Mutchen, USA) at −198° C.

Poly(D,L-lactide-co-glycolide) (PLGA, 110,000 g/mole, 50:50 by moleratio of lactide to glycolide, Resomer RG506) was obtained fromBoehringer Ingelheim (Ingelheim, Germany). Collagraft, collagen-calciumphosphate ceramic graft material, was purchased from Zimmer (Warsaw,Ind., USA). All chemicals were obtained from Sigma Co. (St Louis, Mo.,USA) and used as received unless it is stated otherwise.

B. Methods

(1). Preparation of BSM-PLGA Composite Scaffold and PLGA Scaffold

The BSM-PLGA composite scaffold was fabricated using a solventcasting/particulate leaching process from mixtures of PLGA and BSMpowder. To obtain the desired structural configuration and surfaceproperties of the scaffold, fabrication parameters such as compositionratios of PLGA to BSM and porogen were controlled. As a porogen, sodiumchloride (NaCl) crystals were sieved to obtain a size range of 300 to500 μm to target a pore size range of 100˜250 μm, which is ideal fortissue engineering applications. The particles were added to 20 wt/vol %PLGA in methylene chloride and BSM powder and thoroughly mixed using avortex. PLGA and BSM powder were mixed with equal proportion by weight.The weight ratio of PLGA and BSM to NaCl was 1:10. This formulation wasdetermined by previous experiments that achieved the best structuralintegrity for a bone scaffold. The BSM-PLGA composite scaffoldcontaining NaCl was air-dried for 24 hr and vacuum-dried (5˜10 mbar) for48 hr to remove residual solvent. The scaffolds were immersed indeionized water for 48 hr with water change every 6 hr to remove theembedded salt. The scaffolds were lyophilized and stored in a desiccatoruntil use. PLGA scaffolds without BSM were used as a control, and werefabricated using the same method as the BSM-PLGA composite scaffoldpreparation (1:10).

(2). Scaffold Characterization

Surface and cross-sectional morphology of all scaffolds was examined byscanning electron microscopy (SEM; Model S-2260N, Hitachi Co. Ltd.,Japan). Samples were observed under an environmental SEM (Backscatterelectron (BSE) mode) without any conductive coating. The average poresize and porosity of the composite scaffolds were analyzed by a mercuryintrusion porosimeter (AutoPore IV 9500, Micromeritics Co. Ltd.,Norcross, Ga., USA). The porosimeter was set at a solid penometer volumeof 6.7˜7.3 mL, and a mercury filling pressure of 3.4 KPa and intruded toa maximum pressure of 414 MPa. Individual samples of 0.1 g wereanalyzed.

A fluid uptake test was used to determine the hydrophilicity of thecomposite scaffold. For the evaluation of fluid uptake capacity, thecomposite scaffold was immersed in PBS at pH 7.4 for 30 sec. Percentfluid uptake was calculated according to the following equation:Fluid uptake (%)=(W −W _(o))/W _(o)×100where W_(o) is the dry sample and W_(s) is the wet sample.(3). Mechanical Testing

Mechanical properties were assessed on the cylinder-shaped BSM-PLGAcomposite scaffolds (5 mm of thickness×10 mm of diameter) using amechanical tester with a compression interface diameter of 2 inches(Instron Model 5544, Canton, Mass., USA). A two kN load cell at acrosshead speed of 0.4 mm/min was used for compression testing. Thetesting was performed under both dry and wet conditions. The compressivestrength was determined from the maximum load recorded. Young's moduluswas calculated from the slope of the initial linear segment of thestress-strain curve at maximum stress. At least four specimens weretested for each scaffold, and the average and the standard deviationwere calculated.

(4). Biological Activity Testing

i. Cell Culture

Two different types of cells were used in this study, embryonic stem(ES) cells, which are capable of becoming bone cells, and primary bovineosteoblasts (bOBs). Human ES cells (hES H7) (Thomson J A et al. Science1998;282:1145-7) were maintained on mitomycin C-treated mouse embryonicfibroblasts (MEF, feeder layer) in a growth medium: Knockout highglucose DMEM supplemented with 500 U/mL penicillin, 500 μg/mlstreptomycin, 2 mM GlutaMAX I, 1% non-essential amino acid solution, 0.1mM β-mercaptoethanol, 4 ng/mL bFGF, 1 ng/mL human LIF (Chemicon,Temecula, Calif.), 8% Serum Replacement (SR) and 8% Plasmanate (Bayer).To induce the formation of embryoid bodies (EBs), hES cells weretrypsinized and approximately 1×10⁷ hES cells transferred to a 100 mmPetri dish. Embryoid bodies were grown in the same culture medium as hEScells without LIF, Plasmanate and bFGF. The EBs were cultured for 5 daysand dissociated with trypsin and plated on culture dishes. All reagentsfor cell culture were purchased from Invitrogen Co. (Gibco Cell Culture,Carlsbad, Calif., USA).

Bovine OBs were isolated from the periosteum of calf forelimb. Theperiosteum was harvested from fresh tissue and cut into 1 cm² piecesunder aseptic conditions. The periosteum fragments were placed on 100 mmtissue culture dishes with culture medium (Medium 199 supplemented with10% fetal calf serum (FCS), 50 mg ascorbic acid, 100 U/mL Penicillin and100 μg/mL Streptomycin). The medium was changed every 3 days until aconfluent monolayer of cells was formed. The periosteal tissue wasremoved and the bOBs were subcultured with 0.05% trypsin-EDTA andfurther expanded.

ii. Cytotoxicity Assessment

Scaffolds 3×3×2 mm in size were placed at the center of subconfluenthEBs and bOBs on 96-well plates. Latex fragments of equal size served asa positive control, and cells with media only served as a negativecontrol. The cell-material contact was maintained for 7 days at 37° C.in 5% CO₂ and, culture medium was changed every 3 days. After 7 days,scaffolds were removed and the culture tested for cell viability usingNeutral Red and mitochondrial metabolic activity by3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide (MTT)assay (Pariente J L, Kim B S, Atala A. J Biomed Mater Res 2001;55:33-9).

For the cell viability assay, cells were rinsed with 1×PBS and 0.005wt/vol % Neutral Red in culture medium was added to each well. After 4hr of incubation at 37° C., the Neutral Red solution was removed and dyeextraction was performed by adding 1% acetic acid in 50% ethanolsolution per well. Absorbance was measured using a microplate reader(EL_(x) 800, Bio-Tex Instruments Inc., Winooski, Vt., USA) at 540 nm.The intensity of red color obtained was directly proportional to theviability of cell populations and inversely proportional to thecytotoxicity of the scaffolds.

For the mitochondrial metabolic activity, cells were rinsed with 1×PBSand 1.0 mg/mL of MTT in PBS containing 1.0 g/L glucose was added to thewells. After 4 hr of incubation at 37° C., the MTT solution was removedand insoluble formazan crystals were dissolved in dimethyl sulfoxide(DMSO). Absorbance was measured at 540 nm using a microplate reader. Theintensity of blue color obtained was directly proportional to themetabolic activity of the cell populations and inversely proportional tothe cytotoxicity of the scaffolds.

iii. Cell Adhesion and Proliferation

Biological activity of the BSM-PLGA composite scaffolds was determinedby testing the ability of cells to adhere and proliferate. To assessthese characteristics, both hEBs and bOBs were seeded onto separatecomposite scaffolds. Briefly, all scaffolds were pre-conditioned andsterilized with 70% ethanol for 30 min and rinsed with 1×PBS five times.Scaffolds (8×8×3 mm³) were placed in each well of a 96-well cultureplate. 40 μL of cell suspension (1×10⁶ cells/mL) was seeded on eachscaffold. To observe cell morphology, the cell-scaffold constructs werefixed with 2.5% glutaraldehyde solution. The constructs were dehydratedthrough a series of graded ethanol solutions, followed by observationwith SEM.

Mitochondrial dehydrogenase activity of the cells on scaffolds wasdetermined by MTT assay after 1, 2, 4 and 8 days of culture. In brief,50 μL of 1 mg/mL MTT solution was added to each well containing freshmedium and incubated at 37° C. and 5% CO₂ for 4 hr. The formazan-formedscaffolds were imaged using a digital camera to assess cellulardistribution. The intracellular formazan accumulated in the cytoplasm ofviable cells was solubilized using DMSO. Absorbance was measured at 540nm using a microplate reader. Cell proliferation was assessed by theintensity of blue color obtained, which was directly proportional to themetabolic activity of the cell population. The optical density wasrecorded for quantification (n=5) of adhered cells within the scaffolds.Cell adhesion was further confirmed by scanning electron microscopy ofthe cell seeded scaffolds.

iv. Statistical Analysis

The differences between the composite scaffolds and controls in fluiduptake ability, average pore size, porosity, cell viability andmitochondrial metabolic activity were evaluated by Student's t-test.Data from the cell proliferation assay were analyzed by ANOVA.Differences were considered significant at P<0.05.

2. Fabrication and Characterization of the BSM-PLGA Composite Scaffolds

The BSM-PLGA composite scaffolds, fabricated by the solventcasting/particulate leaching process, exhibited highly porous anduniform interconnected structures. BSM particles with a size rangingfrom 6˜32 μm were homogeneously embedded in the composite scaffolds asdemonstrated by SEM (FIG. 13). FIG. 13 is a cross-sectional observationof (A) the BSM-PLGA composite scaffold, (B) PLGA scaffold and (C)Collagrafts. The BSM-PLGA composite scaffold and PLGA scaffold showsimilar architecture and possess uniform porous structures with aconsistent interconnectivity throughout the entire scaffold. However,Collagrafts shows small pores and minimal interconnectivity. Scale barindicates 500 mm (SEM, original magnification: x100).

There were distinct morphological differences between the BSM-PLGAcomposite scaffold and Collagraft, while the BSM-PLGA composite scaffoldand PLGA scaffold were similar in microstructure (FIG. 13), pore sizeand porosity (Table 1). The exterior surface, serial cross-section andside wall morphologies of the BSM-PLGA composite scaffolds and PLGAscaffolds exhibited a highly porous structure with interconnectivitythat would support adequate cell seeding, adhesion and proliferation.The average pore size of the BSM-PLGA composite scaffolds was121.84±23.44 μm (Table 1). The average pore size and porosity of theBSM-PLGA composite scaffolds were significantly different compared toCollagraft (P<0.05). The BSM-PLGA composite scaffold exhibited uniformlydistributed-interconnecting pores in its inner microstructures whichwere successfully achieved by a particulate leaching process. TABLE 1Physical Properties of the BSM-PLGA Composite Scaffold, PLGA Scaffold,and Collagraft. Average pore Compression Young's Samples size (μm)Porosity (%) strength (MPa) modulus (MPa) BSM-PLGA 121.84 ± 23.44* 94.79 ± 10.76* 0.83 ± 0.05(d)* 5.56 ± 0.64(d)* 0.15 ± 0.02(w)  0.24 ±0.07(w)  PLGA 134.22 ± 12.56* 92.43 ± 3.21* 1.20 ± 0.13(d)* 8.33 ±0.23(d)* 1.14 ± 0.07(w)* 7.69 ± 0.18(w)* Collagraft ® 26.44 ± 7.98 70.69 ± 6.52  0.31 ± 0.04(d)  1.27 ± 0.12(d)  0.12 ± 0.07(w)  0.11 ±0.08(w) *p < 0.05, compared to Collagraft ®.Dry (d) and wet (w) conditions.

The fluid uptake ability of the BSM-PLGA composite scaffolds wassuperior when compared with PLGA scaffold and Collagraft (P<0.01). Thefluid uptake ability of BSM-PLGA composite scaffolds increased with theelevation of the BSM to PLGA ratio due to the hydrophiliccharacteristics of the BSM (data not shown).

Compressive strengths of the BSM-PLGA composite and PLGA scaffolds weresignificantly higher than that of Collagraft in dry condition (Table 1).Under dry condition, the mechanical properties of the composite scaffoldwere similar to PLGA scaffold. On the contrary, the composite scaffoldshowed increased Young's modulus and decreased compressive strengthsunder wet condition due to the water-uptake in BSM. There werestatistical differences between dry and wet conditions (P<0.05) in boththe BSM-PLGA composite scaffold and Collagraft.

3. Biological Activity Testing

(1). Cytotoxicity Assessment

Cytotoxicity assay using Neutral Red indicated that the BSM-PLGAcomposite scaffold, PLGA scaffold and Collagraft did not demonstratesignificant differences in cell viability, when compared to cells grownon tissue culture plates as a negative control (FIG. 14A). Meanwhile,the latex used as a positive control, showed a cell viability of 7.3% ofdissociated cells from hEBs and 10.5% of bOBs when compared to thenegative control, indicating a high cytotoxicity. The mitochondrialmetabolic activity of the BSM-PLGA composite scaffold, PLGA scaffold andCollagraft was also determined by the MTT assay using the direct contactmethod (FIG. 14B). Direct contact with latex significantly decreasedmitochondrial activity by 92.1% in hEBs cultures and 96% in bOBscultures after 7 days. FIG. 14 shows the viability (A) and mitochondrialmetabolic activity (B) of hES cells and bOBs cultured on bone scaffoldsshow similar cell survival in all experimental scaffolds at 7 days byneutral red assay (A) and MTT assay (B). P<0.01 compared to Latex(positive control).

(2) Cell Adhesion and Proliferation

Cross-sectional SEM images of the BSM-PLGA composite scaffold, PLGAscaffold and Collagraft 1 day after seeding with hEBs and bOBs weretaken (data not shown). After 1 day of culture, the cell-seeded BSM-PLGAscaffold showed a homogeneous cell distribution. The PLGA scaffold andCollagraft showed a sparse cell distribution. Cell adhesion to thescaffold was quantified as shown in FIG. 15. There was a significantdifference (P<0.01) between the BSM-PLGA composite scaffold and theothers. A similar observation was achieved with bOBs cells (data notshown). In the BSM-PLGA composite scaffold, the seeded cells wereuniformly distributed and adhered to the walls of the pores. However,there were only a few cells found in the pores of the PLGA scaffold andCollagraft. At 8 days of culture, both cell types proliferated on theBSM-PLGA composite scaffold, indicating that the scaffold was conduciveto cell adhesion and proliferation. A uniform cell distribution wasobserved throughout the entire BSM-PLGA scaffold. MTT staining of theBSM-PLGA composite scaffold showed the uniform penetration of formazancrystals on the surface and vertical cross-sections, indicating adequatecell adhesion and distribution throughout the inner structure. Celladhesion on the BSM-PLGA composite scaffold was significantly higherthan on PLGA scaffold and Collagraft (FIG. 15). The percentage of theseeded hEBs and bOBs adhered to the BSM-PLGA composite scaffold was87.70±5.46% and 82.23±8.77% of the initial cell number, respectively.The levels of cell adhesion showed a significant difference (P<0.01)between the BSM-PLGA composite scaffold and Collagraft.

Both the hEBs and bOBs were cultured on the BSM-PLGA composite scaffold,PLGA scaffold and Collagraft for up to 8 days. The rate of the cellproliferation on the BSM-PLGA composite scaffold was significantlyhigher than that of the PLGA scaffold and Collagraft (P<0.05) (FIG. 16),and did not show significant differences between the cell types atdifferent culture time points. FIG. 16 shows cell proliferation of theBSM-PLGA composite scaffold, PLGA scaffold and Collagrafts up to 8 daysculture, as determined by MTT assay. The cells seeded on BSM-PLGAcomposite scaffolds show a higher proliferation rate when compared tothe other experimental scaffolds (P<0.05). Cells proliferated graduallyin the BSM-PLGA composite scaffold, PLGA scaffold and Collagrafts for 8days of culture. The initial number of seeded cells was 4×10⁴cells/scaffold.

This example demonstrates the successful fabrication of a bio-hybridbone scaffold composed of BSM and PLGA that possess necessarycharacteristics for bone tissue regeneration. The BSM-PLGA compositescaffolds are hydrophilic and possess porous structures with aconsistent interconnectivity throughout the entire scaffold whichresulted in uniform cell seeding, adhesion and proliferation. TheBSM-PLGA composite scaffolds are non-toxic, easily fabricated andprovide structural features that may enhance the formation of bonetissue for therapeutic regeneration.

1. A method of producing an artificial composite tissue constructpermitting coordinated motion, comprising: providing a firstbiocompatable structural matrix having sufficient rigidity to providestructural support; seeding the first matrix with an isolated populationof cells selected from the group consisting of cartilage-forming cells,bone-forming cells and combinations thereof; providing a secondbiocompatable flexible matrix; seeding the flexible matrix with anisolated population of muscle progenitor cells (MPCs); and joining thestructural matrix and the flexible matrix to produce an artificialcomposite tissue construct capable of coordinated motion.
 2. The methodof claim 1, further comprising joining a flexible matrix to two separatestructural matrices to provide a flexible linkage therebetween.
 3. Themethod of claim 1, further comprising joining a flexible matrix to astructural matrix and a natural bone structure to provide a flexiblelinkage therebetween.
 4. The method of claim 1, wherein the structuralmatrix is selected from the group consisting of an electrospunsubstrate, a decellularized substrate, and a synthetic polymersubstrate.
 5. The method of claim 1, wherein the flexible matrix isselected from the group consisting of a decellularized bladder submucosasubstrate and an electrospun substrate.
 6. The method of claim 1,wherein the step of joining the structural matrix and the flexiblematrix comprises using a joining technique selected from the groupconsisting of suturing, heating, and gluing with biological glue.
 7. Themethod of claim 1, wherein the structural matrix and the flexible matrixfurther comprise a biological agent selected from the group consistingof nutrients, growth factors, cytokines, extracellular matrixcomponents, inducers of differentiation, products of secretion,immunomodulators, proteins, antibodies, nucleic acids molecules,carbohydrates, and biologically-active compounds which enhance or allowgrowth of the cellular network or nerve fibers.
 8. The method of claim7, wherein the biological agent is a growth factor selected from thegroup consisting of transforming growth factor-alpha (TGF-α),transforming growth factor-beta (TGF-β), platelet-derived growth factor(PDGF), fibroblast growth factor (FGF), nerve growth factor (NGF), brainderived neurotrophic factor, cartilage derived factor, bone growthfactor (BGF), basic fibroblast growth factor, insulin-like growth factor(IGF), vascular endothelial growth factor (VEGF), granulocyte colonystimulating factor (G-CSF), hepatocyte growth factor, glial neurotrophicgrowth factor (GDNF), stem cell factor (SCF), keratinocyte growth factor(KGF), and skeletal growth factor.
 9. The method of claim 1, wherein therigid tissue layer is selected from the group consisting of a cartilagetissue layer and a bone tissue layer.
 10. The method of claim 1, whereinthe bone-forming cells are selected from the group consisting ofosteogenic cells, osteoblasts, osteocytes, osteoclasts, and bone-liningcells.
 11. The method of claim 1, wherein the cartilage-cells areselected from the group consisting of chondrocytes and chondroblasts.12. The method of claim 1, wherein the first biocompatable structuralmatrix is a composite scaffold.
 13. A method for treating a subject witha limb or digit disorder, comprising: providing an artificial compositetissue construct having coordinated motion comprising a firstbiocompatable structural matrix having sufficient rigidity to providestructural support seeded with an isolated population of cells selectedfrom the group consisting of cartilage-forming cells, bone-forming cellsand combinations thereof, and a second biocompatable flexible matrixseeded with an isolated population of muscle progenitor cells (MPCs),wherein the structural matrix and the flexible matrix are joined;joining the flexible matrix to the structural matrix and a natural bonestructure to provide a flexible linkage therebetween; and monitoring thesubject for an improvement in the limb or digit disorder.
 14. The methodof claim 13, further comprising joining a flexible matrix to twoseparate structural matrices to provide a flexible linkage therebetween.15. The method of claim 13, wherein the step of monitoring theimprovement comprises monitoring parameters selected from the groupconsisting of the mobility of the digit or limb, the extensor and flexorfunction, the motor function, the sensory function, the contractileresponse, and the tensile response.
 16. A biocompatible compositescaffolding system capable of providing structural support forengineered bone tissue comprising a biodegradable synthetic polymer andnaturally derived collagen matrix.
 17. The biocompatible compositescaffolding system of claim 16, wherein the synthetic polymer isselected from the group comprising poly(lactide-co-glycolide) (PLGA),poly(lactide) (PLA), poly(glycolic acid) (PGA), poly(caprolactone),polycarbonates, polyamides, polyanhydrides, polyamino acids, polyorthoesters, polyacetals, polycyanoacrylates, degradable polyurethanes,hydroxyapatite (HA), tricalcium phosphate (TCP), and calcium sulfate.18. The biocompatible composite scaffolding system of claim 16, whereinthe naturally derived collagen matrix is submucosa.
 19. Thebiocompatible composite scaffolding system of claim 18, wherein thesubmucosa is bladder submucosa (BSM).
 20. The biocompatible compositescaffolding system of claim 16, wherein the composite scaffolding systemis fabricated using a solvent casting/particulate leaching process. 21.The biocompatible composite scaffolding system of claim 16, wherein thebiocompatible composite scaffolding system has substantially uniformporous structures having an average pore diameter ranging from about 50to about 250 μM and a porosity greater than about 50%.
 22. Thebiocompatible composite scaffolding system of claim 21, wherein theaverage pore diameter ranges from about 90 to about 150 μm.
 23. Thebiocompatible composite scaffolding system of claim 21, wherein theporosity is greater than about 80%.
 24. A method of fabricating abiocompatible composite scaffolding system capable of providingstructural support for regenerated bone tissue comprising: selecting apore size, obtaining porogens of said selected size; adding the porogensto a solution containing a synthetic polymer and submucosa and mixing todistribute said porogens and form a composite scaffold; drying saidcomposite scaffold to remove residual solvent; and removing saidporogens from said composite scaffold.
 25. The method of claim 24,wherein said porogens are selected from the group consisting of salts,sodium hydroxide, sugars, waxes, gelatins, naphthalene, natural orsynthetic water soluble polymers, natural or synthetic non-water solublepolymers, degradable polymers, non-degradable polymers, partiallydegradable polymers, and mixtures thereof.
 26. The method of claim 25,wherein the salts are selected from the group consisting of sodiumchloride, potassium chloride, sodium fluoride, potassium fluoride,sodium iodide, sodium nitrate, sodium sulfate, sodium iodate, andmixtures thereof.
 27. The method of claim 24, wherein the method furtherincludes sieving porogens through a sieve to obtain said selected sizeporogens.
 28. The method of claim 24, wherein said synthetic polymer isselected from the group comprising poly(lactide-co-glycolide) (PLGA),poly(lactide) (PLA), poly(glycolic acid) (PGA), poly(caprolactone),polycarbonates, polyamides, polyanhydrides, polyamino acids, polyorthoesters, polyacetals, polycyanoacrylates, degradable polyurethanes,hydroxyapatite (HA), tricalcium phosphate (TCP), and calcium sulfate.29. The method of claim 24, wherein the synthetic polymer ispoly(D,L-lactide-co-glycolide) (PLGA) and the submucosa is bladdersubmucosa (BSM).
 30. The method of claim 29, wherein the weight ratio ofPLGA and BSM to NaCl is about 1:10.
 31. The method of claim 24, whereinthe step of removing said porogens from said composite scaffoldcomprises immersion in water.